Fiber-hydrogel composite surgical meshes for tissue repair

ABSTRACT

The presently disclosed composition and methods are provided for a hydrogel or nanofiber-hydrogel composite integrated with a surgical scaffold or mesh. A surgical scaffold device comprised of laminar composite is disclosed for the purpose of reducing foreign body response, managing tissue-materials interface, and improving the integration of the surgical mesh with the surrounding tissue of a subject.

CROSS-REFERENCE TO RELATED APPLICATION

This application is is a national stage application, filed under 35U.S.C. § 371, of International Application No. PCT/US2016/047282, filedon Aug. 17, 2016, which claims the benefit of priority under 35 U.S.C. §119(e) to U.S. Provisional Application No. 62/206,011, filed on Aug. 17,2015 entitled, “Composite Material for Tissue Restoration”. Thisapplication is also related to International Patent ApplicationPCT/US15/45494, filed Aug. 17, 2015 and entitled, “Composite Materialfor Tissue Restoration”. The contents of these related applications areincorporated herein by reference in their entireties.

BACKGROUND 1. Field

The present disclosure relates to composite materials and methods thatrepair soft tissue defects while promoting soft tissue regeneration.

2. Description of Related Art

Soft tissue defects resulting from trauma, oncologic resection, orcongenital malformation are difficult to treat by conventional means.Current therapies, including tissue rearrangements or tissue transfer,cause donor site defects. Other therapies, such as prosthetic implants,lead to fibrosis and encapsulation. Existing strategies to promotetissue ingrowth are also inadequate for the treatment of soft tissuedefects. Current acellular matrices result in flat, fibrotic sheets oftissue rather than the soft, three-dimensional tissue required for idealreconstructions. Finally, while fat grafting can restore soft tissuedefects, its wider use is hampered by variable graft survival andlimited volumes of restoration. An ideal approach to soft tissuereconstruction would encourage regeneration of soft tissues such asadipose tissue or fascia in vivo followed by implantation of the tissuesto promote regeneration. However, adipose tissue or fascial regrowthrequires a suitable matrix for cells to attach, migrate, proliferate,differentiation, and organize into new tissue. Much of the nativeextracellular matrix (ECM) is missing at the repair site. Therefore,recreating a synthetic matrix that not only immediately restores thelost tissue volume, but also reconditions the microenvironment, supportshost cell infiltration, and encourages regeneration of soft tissue,becomes an essential task when repairing soft tissue defects usingadipose or fascial tissue-based reconstruction.

Hydrogels offer several advantages as a material for soft tissuereconstruction. However, to achieve sufficient mechanical property,higher crosslinking densities are usually required. Under theseconditions, however, host tissue cells (e.g., adipocyte progenitors andendothelial progenitors) are not able to penetrate and grow into thescaffolds. In case of degradable hydrogels, scarring and fibrous tissueformation are typical because ingrowth of host tissue occurs too slowly,or at least at a pace slower than the absorption of the fiber material.

Recently, functionalized nanofibers have been developed to serve as ECMmimics to support various cell activities. FDA-compliant syntheticbiodegradable poly-α-esters, such as polycaprolactone (PCL) orpoly(lactide-co-glycolide) (PLGA) can be used to generate nanofibersthrough a process known as electrospinning Biodegradable sutures andimplants prepared from these polymers have been widely used clinicallydue to their excellent track record on biocompatibility. Variousnanofibers of varying diameters and topographies for stem cellengineering applications have been developed. These nanofibers, however,do not offer macroscopic structures, making them difficult to use as 3Dscaffolds.

Given the various problems associated with such conventional methods andsystems, there is still a need in the art for improved solutions tohealing soft tissue defects. The present disclosure provides a solutionfor this need that overcomes the various problems noted in the art.

SUMMARY

The invention is based, at least in part, upon identification ofscaffold complexes having polymeric fiber components that possessimproved properties (e.g., improved qualities for reconstruction of softtissue, as detailed further infra).

In certain aspects, the invention provides a laminar sheet scaffold,composed of nanofiber-hydrogel composite formed around a surgical meshsheet with an interpenetrating structure, thus forming an integratednetwork.

In other aspects, the invention provides a laminar sheet scaffold,composed of nanofiber-hydrogel composite prepared from surfacefunctionalized electrospun fiber sheet and hydrogel, whereas interfacialbonding is introduced between fibers and hydrogel network, thus formingan integrated network.

In additional aspects, the invention provides a laminar sheet scaffold,composed of surgical mesh sheet and hydrogel composite, whereasinterfacial bonding is introduced between surgical mesh surface andhydrogel network, thus forming an integrated network.

In one aspect, the invention provides a scaffold complex that includes apolymeric fiber having a mean diameter of from about 100 nm to about8000 nm covalently linked to a hydrogel material, where the ratio offiber to hydrogel material is from about 1:10 to about 10:1 on acomponent-mass basis, or from about 1 to 50 mg/mL on a concentrationbasis.

In one embodiment, the polymeric fiber includes a biocompatiblebiodegradable polyester. Optionally, the polymeric fiber includespolycaprolactone.

In another embodiment, the hydrogel material is present in the complexin a functional network.

In an additional embodiment, the ratio of fiber to anhydrous hydrogelmaterial is from about 1:10 to about 10:1.

In another embodiment, the polymeric fiber includes a non-wovenpolymeric fiber.

In certain embodiments, the polymeric fiber includes an electrospunpolycaprolactone fiber. Optionally, the polymeric fiber includes asynthetic polymeric material comprising a poly(lactic-co-glycolic acid),a poly(lactic acid), and/or a polycaprolactone, or a combinationthereof.

In one embodiment, the complex is formulated to be substantiallybiocompatible. Optionally, the polymeric fiber includes a biologicalpolymeric material that includes a silk, a collagen, a chitosan, and/ora combination thereof.

In one embodiment, the hydrogel material includes hyaluronic acid.Optionally, the hydrogel material includes a hydrogel material thatincludes a poly(ethylene glycol), a collagen, a dextran, an elastin, analginate, a fibrin, a alginate, a hyaluronic acid, a poly(vinylalcohol), a derivative thereof, or a combination thereof.

In certain embodiments, the hydrogel material includes a processedtissue extracellular matrix.

In one embodiment, the processed tissue extracellular matrix isderivable from an adipose tissue.

In another embodiment, the scaffold complex includes a non-wovenpolycaprolactone fiber.

In one embodiment, the hydrogel material includes a hyaluronic acidsubstantially covering at least a portion of an outer surface of thepolycaprolactone fiber.

In certain embodiments, the hydrogel material is bonded to the outersurface of the polymer fiber.

In another embodiment, the scaffold complex further includes acrosslinking moiety present in an amount effective to introduce bondingbetween polymer fiber and hydrogel material.

In certain embodiments, the scaffold complex includes a plurality ofpores present on or within a surface of the scaffold complex, where thepores are present at a concentration of at least about 50 pores per cm²of the surface, and where at least 80% of the pores have an average porediameter on the surface is at least about 5 microns.

In additional embodiments, the scaffold complex further includes across-linking moiety present in an amount effective to inducecross-linking between polycaprolactone fiber and hyaluronic acid.

Optionally, the scaffold complex promotes tissue growth and cellinfiltration when implanted into a target tissue present in a humansubject.

In certain embodiments, the scaffold complex is substantiallybiodegradable when implanted into a human tissue.

In one embodiment, the scaffold complex is substantiallynon-biodegradable when implanted into a human tissue.

In another embodiment, the scaffold complex further includes atherapeutic agent selected from a cell, a small molecule, a nucleicacid, and a polypeptide.

Another aspect of the invention provides an implantable biomaterial thatincludes a scaffold complex of the invention.

In certain embodiments, the implantable material is substantiallyacellular and/or is substantially free of polypeptides.

In one embodiment, the implantable material is formulated foradministration by injection.

In another embodiment, the implantable material is formulated forsubdermal administration.

An additional aspect of the invention provides a kit containingimplantable material of the invention.

A further aspect of the invention provides a medical device forretaining tissue shape in a subject undergoing a surgical procedure,that includes the scaffold complex and/or the implantable material ofthe invention in an amount effective to provide for the retention of atissue shape when administered to the subject.

Another aspect of the invention provides a method for preparing animplant for tissue or cartilage repair, the method involving the stepsof: providing an acellular, three-dimensional scaffold that includespolymeric fibers oriented to produce a plurality of pores, where atleast a portion of the polymeric fibers are cross-linked to otherpolycaprolactone fibers; disposing a composition that includes ahydrogel material on the polymeric fibers to form a complex; andreacting or stabilizing the complex to form a stabilized implant,thereby preparing the implant.

Optionally, the tissue includes a soft tissue.

A further aspect of the invention provides a method for preparing animplant for tissue or cartilage repair, the method involving the stepsof: providing an acellular, three-dimensional scaffold that includespolymeric fibers oriented to produce a plurality of pores; disposing acomposition that includes a hydrogel material on the polymeric fibers toform a complex; and reacting or stabilizing the complex to form astabilized implant where at least a portion of the polymeric fibers arecross-linked to the hydrogel material.

In certain embodiments, the three-dimensional scaffold includes reactivepolycaprolactone fibers.

A further aspect of the invention provides a method for preparing animplant for tissue or cartilage repair, the method involving the stepsof: providing an acellular, three-dimensional scaffold that includespolymeric fibers oriented to produce a plurality of pores; disposing acomposition that includes a hydrogel material on the polymeric fibers toform a complex; and reacting or stabilizing the complex to form astabilized implant where at least a portion of the polymeric fibers arecross-linked to the hydrogel material.

An additional aspect of the invention provides a method for resolving atissue defect resulting from a trauma or surgical intervention, themethod involving distending the tissue, where distending the tissueincludes implanting an effective amount of the scaffold complex of theinvention into the tissue to thereby distend it.

Another aspect of the invention provides a method for reducing orreversing a tissue defect resulting from an aging-associated disease,disorder or condition, the method involving distending the tissueincluding the tissue, where distending the tissue includes implanting aneffective amount of a scaffold complex of the invention into the tissueto thereby distend it.

Optionally, the tissue defect includes pleural tissue, muscle tissue,skin, or a combination thereof.

In at least one aspect, the invention provides a composite material thatincludes a gel and at least one nanostructure disposed within the gel.The gel can be hydrogel or any other suitable gel. The nanostructure canbe a nanofiber or any other suitable nanostructure. The nanostructurecan be covalently bonded to the gel. The nanostructure can be made ofpolycaprolactone (PCL) or any other suitable material.

In at least another aspect, the invention provides a method for healinga soft tissue defect comprising applying a composite material to a softtissue defect, wherein the composite material includes a gel and ananostructure disposed within the gel.

In still another aspect, the invention provides a method formanufacturing a composite material for use in healing soft tissuedefects w include providing a gel and disposing nanofibers within thegel.

In another aspect, a surgical device is provided that includes i) alaminar scaffold complex including a polymeric fiber sheet (optionally anon-woven and/or electrospun polymeric fiber sheet), having a mean fiberdiameter of from about 100 nm to about 8000 nm operably linked to ahydrogel material; and ii) a surgical mesh material.

Where applicable or not specifically disclaimed, any one of theembodiments described herein are contemplated to be able to combine withany other one or more embodiments, even though the embodiments aredescribed under different aspects of the invention.

These and other embodiments are disclosed or are obvious from andencompassed by, the following Detailed Description.

BRIEF DESCRIPTION OF THE DRAWINGS

The following detailed description, given by way of example, but notintended to limit the invention solely to the specific embodimentsdescribed, may best be understood in conjunction with the accompanyingdrawings.

FIG. 1A is illustrates the structure of an embodiment of a composite inaccordance with this disclosure, showing nanostructures disposed in agel, and in particular, the covalent attachment of the nanostructure tofunctional groups in the gel.

FIG. 1B shows a light microscope image of a fully swollen composite asillustrated in FIG. 1;

FIG. 1C is an image of the macroscopic appearance of a hydratedcomposite as illustrated in FIG. 1;

FIG. 1D shows a scanning electron micrography (SEM) image of adehydrated composite as illustrated in FIG. 1, revealingultra-structural similarity to ECM;

FIG. 2A depicts stress-strain curves of an embodiment of the compositeof FIG. 1 plotted against HA Hydrogel alone, revealing improved elasticmodulus compared to hydrogel at the same crosslinking density;

FIG. 2B depicts a fatigue test showing that the embodiment of acomposite of FIG. 2A retains similar degree of robustness of mechanicalintegrity compared to regular hydrogel;

FIGS. 3A and 3B show fluorescence and overlay (FIG. 3A) with phasecontrast images (FIG. 3B) of ASCs cultured in nanofiber-HA hydrogelcomposite for 4 days;

FIGS. 3C and 3D show fluorescence and overlay (FIG. 3C) with phasecontrast images (FIG. 3D) of ASCs cultured in regular HA hydrogel for 4days;

FIGS. 4A and 4B show a fluorescence image and overlay (FIG. 4A) withphase contrast image (FIG. 4B) contrasting ASCs migrating from spheroidsalong aligned 650-nm nanofibers.

FIG. 5A is a photograph showing appearance of nanofiber-hydrogelcomposite in situ under rat inguinal fat pad;

FIG. 5B shows H&E staining images of sections from tissues around thecomposite harvested at 2 weeks after implantation; and

FIG. 5C shows H& E staining images of tissue sections collected fromcomposite-tissue interface at 4 weeks, showing cell infiltration.

FIG. 6A depicts a synthesis scheme for the polycaprolactone (PCL)fiber-HA hydrogel composite.

FIG. 6B depicts a schematic illustration of the composite structure withinterfacial bonding between PCL fibers and HA chain network.

FIG. 6C depicts optical images showing the general appearance of afreshly prepared, cylindrical fiber-HA hydrogel composite (left) and aHA hydrogel (right) with the same dimensions (scale bar=5 mm).

FIG. 6D depicts optical images of the same set of samples afterlyophilization and rehydration.

FIG. 6E depicts SEM images of cross-section of an HA hydrogel (scalebar=40 μm).

FIG. 6F depicts SEM images of cross-section of PCL fiber-HA hydrogelcomposite (scale bar=100 μm).

FIG. 6G depicts SEM images of cross-section of decellularized native fattissue (scale bar=10 μm).

FIG. 7A depicts the effect of fiber diameter and the interfacial bondingon reinforcing compressive modulus of HA hydrogel. HA hydrogel andcomposites were prepared based on 4.5 mg/ml of HA. The values of stresswere measured at 50% of strain. *p<0.05 (Student-t test).

FIG. 7B depicts the effect of fiber diameter and the interfacial bondingon reinforcing compressive modulus of PEG hydrogel. PEG hydrogel andcomposites were prepared based on 30 mg/ml of PEGSH and 20 mg/ml ofPEGDA, and 1.0-μm PCL fibers were used to synthesize the fiber-PEGhydrogel composites. The values of stress were measured at 50% ofstrain. *p<0.05 (Student-t test).

FIG. 8A depicts the effect of interfacial bonding density and the fiberdiameter on reinforcing shear storage modulus of HA hydrogel. *p<0.05(Student-t test).

FIG. 8B depicts the effect of interfacial bonding density and the fiberdiameter on reinforcing shear storage modulus of PEG hydrogel. Thevalues of shear storage modulus were measured at 1-Hz frequency. *p<0.05(Student-t test).

FIG. 8C depicts the effect of interfacial bonding density and the fiberdiameter on reinforcing shear storage modulus of HA hydrogel. The valuesof shear storage modulus were measured at 1-Hz frequency. *p<0.05(Student-t test).

FIG. 8D depicts the effect of interfacial bonding density and the fiberdiameter on reinforcing shear storage modulus of HA hydrogel. The valuesof shear storage modulus were measured at 1-Hz frequency. *p<0.05(Student-t test).

FIG. 9A depicts the effect of fiber-loading amount on shear storagemodulus of HA hydrogel. The HA hydrogel and composites were synthesizedusing a 10-mg/ml of HA. Shear storage moduli are measured at 1-Hzfrequency. Blue arrows indicate conditions for both composites with a 1to 2 of molar ratio of SH groups to (DA+MAL) groups. *p<0.05 (Student-ttest).

FIG. 9B depicts the effect of fiber-loading amount on shear storagemodulus of HA hydrogel. The HA hydrogel and composites were synthesizedusing 4.5-mg/ml of HA. Shear storage moduli are measured at 1-Hzfrequency. Blue arrows indicate conditions for both composites with a 1to 2 of molar ratio of SH groups to (DA+MAL) groups. *p<0.05 (Student-ttest).

FIG. 10A depicts the mechanical strength of the fiber-HA hydrogelcomposite under different frequencies. Shear storage modulus of the HAhydrogel and the composites is measured against different frequencies ofshear loading.

FIG. 10B depicts the mechanical strength of the fiber-HA hydrogelcomposite under different rehydration. Comparison for the compressivestress of the composites before and after rehydration (strain=40%).

FIG. 10C depicts the mechanical strength of the fiber-HA hydrogelcomposite under different cyclic loading. Compressive stresses of an HAhydrogel and the corresponding composite are measured against cyclicloading (strain=25%).

FIG. 11A depicts the migration ability of human adipose-derived stemcells (hASCs) in HA hydrogel on Day 27. The HA hydrogel control and thetwo composites were selected to exhibit similar compressive moduli ofaround 1.9 kPa. F-actin and nuclei of hASCs were stained with AlexaFluor® 568 phalloidin (red) and DAPI (blue), respectively. Nanofiberswere labeled with Alexa Fluor® 647 (white). Scale bars=100 μm.

FIG. 11B depicts the migration ability of human adipose-derived stemcells (hASCs) in nanofibers-HA hydrogel composite on Day 27. The HAhydrogel control and the two composites were selected to exhibit similarcompressive moduli of around 1.9 kPa. F-actin and nuclei of hASCs werestained with Alexa Fluor® 568 phalloidin (red) and DAPI (blue),respectively. Nanofibers were labeled with Alexa Fluor® 647 (white).Scale bars=100 μm.

FIG. 11C depicts the migration ability of human adipose-derived stemcells (hASCs) in RGD-nanofibers-HA hydrogel composite on Day 27. The HAhydrogel control and the two composites were selected to exhibit similarcompressive moduli of around 1.9 kPa. F-actin and nuclei of hASCs werestained with Alexa Fluor® 568 phalloidin (red) and DAPI (blue),respectively. Nanofibers were labeled with Alexa Fluor® 647 (white).Scale bars=100 μm.

FIG. 11D depicts the migration ability of human adipose-derived stemcells (hASCs) in RGD-nanofibers-HA hydrogel composite on Day 27. The HAhydrogel control and the two composites were selected to exhibit similarcompressive moduli of around 1.9 kPa. Yellow arrows in (d) and (e)indicate cells adhering to fibers or fibers clusters. F-actin and nucleiof hASCs were stained with Alexa Fluor® 568 phalloidin (red) and DAPI(blue), respectively. Nanofibers were labeled with Alexa Fluor® 647(white). Scale bars=20 μm.

FIG. 11E depicts the migration ability of human adipose-derived stemcells (hASCs) in nanofibers-HA hydrogel composite on Day 27. The HAhydrogel control and the two composites were selected to exhibit similarcompressive moduli of around 1.9 kPa. Yellow arrows in (d) and (e)indicate cells adhering to fibers or fibers clusters. F-actin and nucleiof hASCs were stained with Alexa Fluor® 568 phalloidin (red) and DAPI(blue), respectively. Nanofibers were labeled with Alexa Fluor® 647(white). Scale bars=20 μm.

FIG. 11F depicts the migration ability of human adipose-derived stemcells (hASCs). Schematic illustration of hASCs spheroids in thecomposite structure with interfacial bonding between PCL fibers and HAchain network is shown.

FIG. 12A depicts tissue regeneration mediated by the implanted fiber-HAhydrogel composite and HA hydrogel in 30 days. Macroscopic images of thecomposite before (insets) and after implantation under the inguinal fatpad (scale bar=2 mm) are shown. White stars indicate the implantedmatrices.

FIG. 12B depicts tissue regeneration mediated by the implanted fiber-HAhydrogel composite and HA hydrogel in 30 days. Macroscopic images of theHA hydrogel before (insets) and after implantation under the inguinalfat pad (scale bar=2 mm) are shown. White stars indicate the implantedmatrices.

FIG. 12C depicts tissue regeneration mediated by the implanted fiber-HAhydrogel composite and HA hydrogel in 30 days. H&E and Masson'strichrome stained-images of (i) native fat tissue, (ii) healed tissueafter sham surgery, (iii, v) the fiber-HA hydrogel implanted tissue, and(iv, vi) the HA hydrogel implanted tissue on Day 14 and Day 30 areshown. In the images, H=HA hydrogel, C=fiber-HA hydrogel composite,B=brown adipose tissue, yellow arrow=blood vessel. Scale bar=200 μm.

FIG. 12D depicts tissue regeneration mediated by the implanted fiber-HAhydrogel composite and HA hydrogel in 30 days. H&E and Masson'strichrome stained-images of (i) native fat tissue, (ii) healed tissueafter sham surgery, (iii, v) the fiber-HA hydrogel implanted tissue, and(iv, vi) the HA hydrogel implanted tissue on Day 14 and Day 30 areshown. Blue staining from Masson's trichromatic staining indicates totalcollagen in examined tissue. In the images, H=HA hydrogel, C=fiber-HAhydrogel composite, B=brown adipose tissue, yellow arrow=blood vessel.Scale bar=200 μm.

FIG. 13A depicts a schematic diagram of preparing surface-modifiedfibers with MAL via PAA-grafting method.

FIG. 13B depicts average densities of carboxyl groups on fibers afterthe PAA-grafting with 3 and 10% (v/v) of acrylic acid (*p<0.05, n=6).

FIG. 14 depicts shear storage moduli of HA hydrogel with various molarratios of SH to DA prepared with 4.5 mg/ml HA-SH.

FIG. 15A depicts shear storage moduli of fiber-HA hydrogel compositesprepared from various amount of fibers. The average diameter of fibersis 686 nm, MAL surface density on the fibers was 100 nmol/mg, and thecomposites were prepared with 4.5 mg/ml of HA-SH and 5 mg/ml of PEGDA.Blue arrows indicate 1 to 2 of molar ratio of SH groups to (DA+MAL)groups. *p<0.05 (n=3).

FIG. 15B depicts shear storage moduli of fiber-PEG hydrogel compositeswith various amounts of loaded fibers. *p<0.05 (n=3).

FIG. 16 depicts the average pore size of HA hydrogel and nanofiber-HAhydrogel composite were estimated based on the SEM images of theircross-section (*p<0.05).

FIG. 17A depicts cell infiltration and tissue in-growth through thefiber-HA hydrogel composite on Day 14. The sectioned tissues werestained by H&E for total collagen (blue). Labels: C=fiber-HA hydrogelcomposite, yellow arrow=blood vessel. Scale bar=50 μm.

FIG. 17B depicts cell infiltration and tissue in-growth through thefiber-HA hydrogel composite on Day 14. The sectioned tissues werestained by Masson's Trichrome for total collagen (blue). Labels:C=fiber-HA hydrogel composite, yellow arrow=blood vessel. Scale bar=50μm.

FIG. 17C depicts cell infiltration and tissue in-growth through thefiber-HA hydrogel composite on Day 30. The sectioned tissues werestained by H&E for total collagen (blue). Labels: C=fiber-HA hydrogelcomposite, yellow arrow=blood vessel. Scale bar=50 μm.

FIG. 17D depicts cell infiltration and tissue in-growth through thefiber-HA hydrogel composite on Day 30. The sectioned tissues werestained by Masson's Trichrome for total collagen (blue). Labels:C=fiber-HA hydrogel composite, yellow arrow=blood vessel. Scale bar=50μm.

FIG. 18 depicts SEM images of cross-section of the decellularized fattissue (upper panel) and the fiber-HA hydrogel composite (lower panel).

FIG. 19A depicts migration ability of hASCs in HA hydrogels (G′=24.85μ2.92 Pa) on Day 4. The HA hydrogel was fabricated with 2.5 mg/ml ofHA-SH and 5.0 mg/ml of PEGDA. Scale bar=100 μm.

FIG. 19B depicts migration ability of hASCs in 1.0-μm fiber-HA hydrogelcomposite (G′=32.29μ 2.16 Pa) on Day 4. The composites were fabricatedwith 2.5 mg/ml of HA, 5.0 mg/ml of PEGDA and 10 mg/ml fibers. Scalebar=100 μm.

FIG. 19C depicts migration ability of hASCs in 286-nm fiber-HA hydrogelcomposite (G′ 39.56μ 1.26 Pa) on Day 4. The composites were fabricatedwith 2.5 mg/ml of HA, 5.0 mg/ml of PEGDA and 10 mg/ml fibers. Scalebar=100 μm.

FIG. 20A depicts injectable formulation. The fiber-hydrogel compositecan be formulated for injectable applications.

FIG. 20B depicts the injectable composite is stable immediately afterinjection.

FIG. 20C depicts the injectable composite remains non-dispersive inwater with shape and volume retention.

FIG. 20D depicts cell infiltration and tissue in-growth through theinjectable fiber-HA hydrogel composite on Day 30, showing extensivecellular remodeling and adipocyte formation. The sectioned tissues werestained by H&E. Labels: c=fiber-HA hydrogel composite.

FIG. 21 depicts the structure and preparation scheme of a compositesurgical scaffold device comprised of a fibrous surgical mesh andnanofiber-hydrogel composite. Insert at the lower right shows an opticalimage of a composite mesh lyophilized and rehydrated back to theoriginal hydrated state.

FIG. 22 depicts the structure and preparation scheme of compositenanofiber-hydrogel sheet scaffold. Surface functionalize nanofiber sheetis used without additional microfiber mesh. The random or alignednanofiber sheet is functionalized using the same method as described inPCT/US15/45494 without the cryomilling step. The intact sheet is thenused to form composite sheet scaffold.

FIG. 23 depicts optical images of composite nanofiber sheet inlyophilized form (left) and rehydrated form (right).

FIG. 24 depicts the structure and preparation scheme of compositemicrofiber-hydrogel sheet scaffold. Surface functionalized microfibersheet is used without additional nanofiber mesh. The knitted or woven ornonwoven microfiber mesh sheet is functionalized using the same methodas described in PCT/US15/45494. The microfiber sheet is then used toform composite sheet scaffold.

FIG. 25 depicts an optical image of a composite microfiber sheet in therehydrated form. This scaffold configuration was prepared without thenanofiber component. The fiberless HA hydrogel was cast around themicrofiber mesh.

FIGS. 26A and 26B depict the placement of the surgical mesh along theabdominal wall in the rat in vivo model. A drawing depicting surgicalprocedure (FIG. 26A) and an image of mesh implantation in rodent (FIG.26B) are specifically presented.

FIGS. 27A and 27B depict the histologic evaluation of sham surgerycontrol (top row), uncoated polypropylene mesh (middle row), andcomposite mesh (bottom row) at day 4 following in vivo implantation.Representative Hematoxylin and Eosin (H&E; FIG. 27A) and Masson'sTrichrome-stained (FIG. 27B) consecutive tissue cross sections wereimaged at 5× magnification (left column) and 20× magnifications (middleand right columns). Increased cellular infiltration was observed in thecomposite mesh groups when compared to control and mesh only groups.

DETAILED DESCRIPTION

The present invention relates to composite materials comprising ahydrogel and a nanostructure for use in methods for reconstruction ofsoft tissue. The invention also relates to methods for repairing orreconstructing a soft tissue injury using a composition comprising ahydrogel and a nanostructure disposed therein. The invention in otheraspects also relates to a method of fabricating a composition for use insoft tissue reconstruction where the composition comprises a hydrogeland a nanostructure disposed therein.

One advantage of the composite sheet scaffolds provided herein is thetendency of these scaffolds to reduce foreign body response, reduceinflammation, and improve tissue-materials interface, ultimatelyimproving the integration of surgical mesh with the surrounding tissueof a subject. The scaffolds described herein can be used for variousapplications, as described infra and as otherwise known or appreciatedin the art.

At least three distinct configurations of the instant composite sheetscaffolds are provided herein. In a first configuration, comprisingsurgical mesh, hydrogel and nanofiber, the configuration is designed tomaximize the ability of the scaffold to encourage cell infiltration fromsurrounding tissue and improve integration of the implant and thetissue, and the full advantages of the nanofiber-hydrogel composite asdefined in parent patent application PCT/US15/45494 are realized. In analternative configuration, which is a simpler configuration comprisingonly the surgical mesh and hydrogel (no nanofiber). This alternativeconfiguration still retains the good mechanical property of a microfibersurgical mesh. The final configuration is a laminar matrix of nanofiberslinked to the hydrogel but without a separate surgical mesh. Thisconfiguration does not have the great tensile strength ofmesh-nanofiber-hydrogel composition, but it is simpler to produce andwill induce less scar formation due to the resorbability of thenanofiber and hydrogel components compared to the permanent surgicalmesh. Thus, this configuration will be ideal in situations such as duralor pleural repair where high tensile strength is not required butminimal scar formation is necessary. Importantly, an integrated networkstructure is the common feature of all exemplified configurations.

The following is a detailed description of the invention provided to aidthose skilled in the art in practicing the present invention. Those ofordinary skill in the art may make modifications and variations in theembodiments described herein without departing from the spirit or scopeof the present invention. Unless otherwise defined, all technical andscientific terms used herein have the same meaning as commonlyunderstood by one of ordinary skill in the art to which this inventionbelongs. The terminology used in the description of the invention hereinis for describing particular embodiments only and is not intended to belimiting of the invention. All publications, patent applications,patents, figures and other references mentioned herein are expresslyincorporated by reference in their entirety.

Although any methods and materials similar or equivalent to thosedescribed herein can also be used in the practice or testing of thepresent invention, the preferred methods and materials are nowdescribed. All publications mentioned herein are incorporated herein byreference to disclose and described the methods and/or materials inconnection with which the publications are cited.

Unless defined otherwise, all technical and scientific terms used hereinhave the meaning commonly understood by a person skilled in the art towhich this invention belongs. The following references, the entiredisclosures of which are incorporated herein by reference, provide oneof skill with a general definition of many of the terms (unless definedotherwise herein) used in this invention: Singleton et al., Dictionaryof Microbiology and Molecular Biology (2nd ed. 1994); The CambridgeDictionary of Science and Technology (Walker ed., 1988); The Glossary ofGenetics, 5th Ed., R. Rieger et al. (eds.), Springer Verlag (1991); andHale & Marham, the Harper Collins Dictionary of Biology (1991).Generally, the procedures of molecular biology methods described orinherent herein and the like are common methods used in the art. Suchstandard techniques can be found in reference manuals such as forexample Sambrook et al., (2000, Molecular Cloning—A Laboratory Manual,Third Edition, Cold Spring Harbor Laboratories); and Ausubel et al.,(1994, Current Protocols in Molecular Biology, John Wiley & Sons,New-York).

The following terms may have meanings ascribed to them below, unlessspecified otherwise. However, it should be understood that othermeanings that are known or understood by those having ordinary skill inthe art are also possible, and within the scope of the presentinvention. All publications, patent applications, patents, and otherreferences mentioned herein are incorporated by reference in theirentirety. In the case of conflict, the present specification, includingdefinitions, will control. In addition, the materials, methods, andexamples are illustrative only and not intended to be limiting.

Definitions

As used herein, a “scaffold complex” includes any covalent associationof two components: a polymeric fiber and a hydrogel material. Thescaffold complex contains the polymeric fiber and hydrogel material in a“functional network”, meaning that the interactions between componentsresults in a chemical, biochemical, biophysical, physical, orphysiological benefit. In addition, a functional network may includeadditional components, including cells, biological materials (e.g.,polypeptides, nucleic acids, lipids, carbohydrates), therapeuticcompounds, synthetic molecules, and the like. In certain embodiments,the scaffold complex promotes tissue growth and cell infiltration whenimplanted into a target tissue present in a human subject.

As used herein, the term “hydrogel” is a type of “gel,” and refers to awater-swellable polymeric matrix, consisting of a three-dimensionalnetwork of macromolecules (e.g., hydrophilic polymers, hydrophobicpolymers, blends thereof) held together by covalent or non-covalentcrosslinks that can absorb a substantial amount of water (e.g., 50%, 60%70%, 80%, 90%, 95%, 96%, 97%, 98%, 99% or greater than 99% per unit ofnon-water molecule) to form an elastic gel. The polymeric matrix may beformed of any suitable synthetic or naturally occurring polymermaterial. As used herein, the term “gel” refers to a solidthree-dimensional network that spans the volume of a liquid medium andensnares it through surface tension effects. This internal networkstructure may result from physical bonds (physical gels) or chemicalbonds (chemical gels), as well as crystallites or other junctions thatremain intact within the extending fluid. Virtually any fluid can beused as an extender including water (hydrogels), oil, and air (aerogel).Both by weight and volume, gels are mostly fluid in composition and thusexhibit densities similar to those of their constituent liquids. Ahydrogel is a type of gel that uses water as a liquid medium.

The definitions of “hydrophobic” and “hydrophilic” polymers are based onthe amount of water vapor absorbed by polymers at 100% relativehumidity. According to this classification, hydrophobic polymers absorbonly up to 1% water at 100% relative humidity (“rh”), while moderatelyhydrophilic polymers absorb 1-10% water, hydrophilic polymers arecapable of absorbing more than 10% of water, and hygroscopic polymersabsorb more than 20% of water. A “water-swellable” polymer is one thatabsorbs an amount of water greater than at least 50% of its own weight,upon immersion in an aqueous medium.

The term “crosslinked” herein refers to a composition containingintramolecular and/or intermolecular crosslinks, whether arising throughcovalent or noncovalent bonding, and may be direct or include across-linker. “Noncovalent” bonding includes both hydrogen bonding andelectrostatic (ionic) bonding.

The term “polymer” includes linear and branched polymer structures, andalso encompasses crosslinked polymers as well as copolymers (which mayor may not be crosslinked), thus including block copolymers, alternatingcopolymers, random copolymers, and the like. Those compounds referred toherein as “oligomers” are polymers having a molecular weight below about1000 Da, preferably below about 800 Da. Polymers and oligomers may benaturally occurring or obtained from synthetic sources.

In certain embodiments, a surgical mesh is employed. In someembodiments, “surgical mesh” refers to a loosely woven sheet which isused as either a permanent or temporary support for organs and othertissues during surgery. Surgical mesh can be created from inorganicand/or biological materials, and can be used in a variety of surgeries.

Soft Tissue Reconstruction

Devastating soft tissue losses from tumor extirpation, trauma, aging, orcongenital malformation affect millions of people each year. The loss oftissues including skin, fat, and muscle lead to major functional andaesthetic disturbances that are difficult to treat by conventionalmeans. As an example, over 300,000 partial mastectomies are performed inthe United States each year, leading to disfiguring breast scars fromthe loss of breast soft tissue. Existing options for soft tissuerestoration have significant drawbacks. Autologous tissue flaps requiresmoving soft tissues from another part of the body in lengthy surgicalprocedures that leave donor-site deficits LoTempio 2010. Plastic andReconstructive Surgery, 126(2), 393-401; Patel 2012. Annals of PlasticSurgery, 69(2), 139-1441. Prosthetic implants are prone to foreign-bodyresponse leading to fibrosis and encapsulation {Calobrace 2014 Plasticand Reconstructive Surgery, 134(1 Suppl), 6S-11; Tsoi 2014. Plastic andReconstructive Surgery, 133(2), 234-249}. Fat grafting involvingplacement of adipocytes harvested during liposuction is limited to smallvolumes and is hampered by poor graft survival {Kakagia 2014 SurgicalInnovation, 21(3), 327-336; Largo 2014 British Journal of PlasticSurgery, 67(4), 437-448}. Finally, injectable hydrogel soft tissuefillers can be used, but these are suitable only for smaller defects andthe volume restoration they provide is transient {Young 2011. ActaBiomaterialia, 7(3), 1040-1049; Varma 2014 Acta Biomaterialia, 10(12),4996-5004}. A new generation of tissue engineering solutions has beenproposed to focus on using hydrogel scaffolds as templates to regeneratesoft tissues such as adipose tissue at the site of reconstruction.

Current Tissue Engineering Approaches to Soft Tissue Reconstruction

Adipose-derived stem cells (ASCs) are a type of mesenchymal stem cell(MSC) that have been identified in wound beds surrounding soft tissuedefects {Salibian 2013 Archives of plastic surgery 40.6: 666-675}. Theycan be differentiated into soft tissues such as fat, when supported witha suitable matrix microenvironment. Other MSCs can migrate into suitablemicroenvironments to reconstitute soft tissues such as fascia, dura, andpleura. Therefore strategies to fill the repair site with functionalmaterials have the potential to enable the regeneration of new tissueusing the endogenous MSCs. Hydrogels have been widely studied as ascaffold matrix for the regeneration of tissue defects due to theirthree-dimensional (3D) nature and elastic properties, which are similarto those of soft tissues. Various methods have been used to generatehydrogel scaffolds with moduli similar to that of native fat tissues (˜2kPa) {Alkhouli 2013 American Journal of Physiology. Endocrinology andMetabolism, 305(12), E1427-35; Sommer 2013 Acta biomaterialia 9.11(2013): 9036-9048} while maintaining their volume and shape againstphysical stress from the surrounding tissue. This requires highercrosslinking density and smaller average pore size {Ryu 2011Biomacromolecules 12.7 (2011): 2653-2659; Khetan 2013 Nature Materials,12(5), 458-465; Li 2014 Journal of Neurotrauma, 31(16), 1431-1438},leading to low cellular infiltration and poor regeneration. The abilityfor hydrogel scaffolds to promote cellular infiltration is the key tosuccessful soft tissue restoration. Lack of vascular infiltration isresponsible for the failure of large-volume fat grafting and othertissue engineering attempts. No currently available materials can fillthe volume lost in soft tissue defects while promoting earlyvascularization and ASC differentiation to regenerate soft tissue.

Hydrogel Matrix

Over the past few years, Li and Wen have developed a hyaluronic acid(HA) hydrogel conjugated with laminin-derived loop peptide(CCRRIKVAVWLC, 10 μM) with optimized pore size and modulus (10-100 Pa)for stem cell transplantation. They have shown that this hydrogelsupports robust neural stem cell (NSC) migration and neurite sproutingfrom the differentiated cells {Li 2014 Journal of Neurotrauma, 31(16),1431-1438}. In a rat controlled cortical injury (CCI) model fortraumatic brain injury, this hydrogel, when injected on day 3 after theCCI injury, promoted significant vasculature network formation fillingthe lesion site (>10 mm) at 4 weeks to 6 months post implantation. Thisimproved angiogenesis was attributed to the ability of this hydrogel toretain and present tissue-secreted growth factors, particularly vascularendothelial growth factor (VEGF). Literature reports also revealed thatsmall HA degradation fragments of 3-10 disaccharide units were potentregulators of endothelial cell proliferation, migration, tubuleformation, and angiogenesis {Slevin 2002 Journal of BiologicalChemistry, 277(43), 41046-41059}. In a recent study, the effectivenessof this HA hydrogel to deliver human fetal tissue derived-NSC spheroidsin a brain lesion site after CCI injury was tested. This HA hydrogeldelivered robust vascular formation inside the scaffold matrix followingtransplantation. Regenerated blood vessels grew into the lesion andpenetrated through the implanted matrix, and supported the survival andgrowth the neuronal progenitors. Even though these studies are not foradipose tissue regeneration, these results confirmed the unique abilityof this optimized HA hydrogel composition in promoting host vascularingrowth. More importantly, the hydrogel matrix is sufficiently porousto allow robust cell migration inside the hydrogel matrix. However,using this HA hydrogel directly for soft tissue reconstruction is notfeasible, as its mechanical property is not sufficiently high tomaintain the integrity of the implantation site—the surrounding adiposetissue has a modulus of more than 10-times higher. Increasingcrosslinking density to improve its modulus will make it poorlypermeable for cell infiltration and migration. A new strategy is neededto increase the mechanical property without significantly decreasing theaverage pore size of the bulk hydrogel. Provided are hydrogel materialsthat contain and/or are isolated from a processed tissue extracellularmatrix, such as extracellular matrix derived and/or derivable from anadipose tissue.

Scaffold Complexes.

Provided herein are scaffold complexes suitable for use medical devicesthat are incorporated into a tissue of a human subject to whom thecomplexes are administered, e.g., by injection or implantation. Thescaffold complexes contain a polymeric fiber, generally having a meandiameter of from about 10 nm to about 10,000 nm, such as about 100 nm toabout 8000 nm, or about 150 nm to about 5,000 nm, or about 100, 150,200, 250, 300, 350, 400, 450, 500, 600, 700, 800, 900, 1,000, 1,500,2,000, 2,500, 3,000, 3,500, 4,000, 4,500, 5,000, 5,500, 6,000, 6,500,7,000, 7,500, or 8,000. As provided herein, the ratio of polymeric fiberto hydrogel material can be determined my any means known in the art.For example, the ratio of polymeric fiber to hydrogel material is fromabout 1:100 to about 100:1 on a component-mass basis, such as about 1:50to about 50:1, or 1:10 to about 10:1, such as 1:5 to about 5:1, such asabout 1:3 to about 3:1. The ratio of polymeric fiber to hydrogelmaterial is also provided as a concentration basis, e.g., a given weightof polymeric fiber per volume of hydrogel material. For example theconcentration is from about 1 to 50 mg/mL. The hydrogel material isgenerally disposed on the polymer fiber, such as being bonded to theouter surface (or an outer surface, depending upon the composition andshape) of the polymer fiber. The scaffold complex is not generally auniform solid material. Instead, scaffold complexes contain a pluralityof pores present on or within a surface of the scaffold complex. Thepresence, size, distribution, frequency and other parameters of thepores can be modulated during the creation of the scaffold complex. Poresize can be from below about 1 micron to up to 100 microns, including 1,2, 3, 4 5, 10, 15, 20, 30, 40, 50, 60 70, 80, 90 or 100 microns, and thesize thereof may be narrowly tailored, e.g., such that at least 40%,such as 50%, 60%, 70%, 80%, 90%, 95% or greater than 95% of the poresare in a desired size or within a desired size range.

The scaffold complexes of the invention are suitable for incorporationinto a tissue of a human subject, and thus they are generally“biocompatible”, meaning capable of interacting with a biological system(such as found in a human subject) without inducing a pathophysiologicalresponse therein and/or thereby. In some embodiments the scaffoldcomplex is provided in order to be durably retained in the tissue.Alternatively, the scaffold complexes are transiently retained in thehuman subject, and are provided as substantially biodegradable.Preferably, a polymeric fiber contains a biocompatible biodegradablepolyester. In a preferred embodiment, the polymeric fiber containspolycaprolactone.

As provided herein, one preferred form of interaction of the complexcontaining polymer fiber and hydrogel includes a crosslinking moiety,generally present in an amount effective to introduce bonding betweenpolymer fiber and hydrogel material, e.g., to induce cross-linkingbetween polycaprolactone fiber and hyaluronic acid.

Scaffold Design for Soft Tissue Restoration

The composite concept has been widely used as a material-reinforcementmechanism. For example, adding hydroxyapatite particles into hydrogelcan increase its stiffness {Wu 2008 Materials Chemistry and Physics107.2 (2008): 364-369}, and the composite tensile modulus increases evenmore for elongated particles {Yusong 2007 Journal of Materials Science,42(13), 5129-5134}. Electrospun nanofiber meshes have been used widelyas a tissue engineering substrate due to their topographical similarityto the native ECM. Of particular interest, the decellularized ECM ofadipose tissue is highly fibrous and porous in nature (FIG. 6G) {Young2011. Acta Biomaterialia, 7(3), 1040-1049}. Several recent studies haveattempted to recapitulate the fibrous components by introducingfragmented poly(lactide) (PLA) or chitosan fibers to a polyethyleneglycol (PEG), polyacrylamide, or alginate hydrogel {Coburn 2011 SmartStructures and Systems, 7(3), 213; #37; Zhou 2011 Colloids and SurfacesB: Biointerfaces, 84(1), 155-162; Shin 2015 Journal of MaterialsChemistry}. The fragmented fibers are mixed with hydrogel precursorsolutions and incorporated into hydrogel during the gelation process tocreate a 3D architecture. These fiber-embedded hydrogels have shownimproved mechanical properties over the corresponding hydrogels.However, there has been no report on testing host cell infiltration invivo. In addition, these hydrogels are non-degradable and requireadhesive ligands for adipocyte adhesion and differentiation.

Nanofiber-Hydrogel Composite Design

To achieve fiber-reinforcement effect while maintaining high porosity inthe hydrogel phase, an electrospun fiber-hydrogel composite that offerssuperior properties as compared to other scaffolds is provided. Beyondblending nanofibers and a hydrogel matrix, which has been reportedpreviously {Coburn 2011 Smart Structures and Systems, 7(3), 213},introduced here are interfacial bonding between fiber surfaces and thehydrogel crosslinking network (FIG. 6). Such a composite design not onlyallows stronger mechanical reinforcement from the solid fiber component,but also allows independent tuning of bulk mechanical properties and theaverage pore size/porosity of the hydrogel phase, enabling both optimalcell infiltration properties and structural integrity. It is furthercontemplated that fibers can be employed as preferred cell adhesionsubstrates for ASCs and endothelial progenitors, therefore acting as aguide to support cell migration and ASC differentiation.

Innovation

In certain aspects, a key innovation is the nanofiber-hydrogel compositedesign with interfacial bonding between nanofiber surfaces and thehydrogel network (FIG. 6A). This engineered composite has the potentialto drastically improve the mechanical property of the hydrogel withoutsignificantly decreasing the average pore size in the hydrogel phase.The introduction of interfacial bonding can offer superior mechanicalstrengthening effect comparing to just physical blending of the twocomponents. This study will map out the range of mechanical properties(compression and shear moduli) attainable with electrospunpolycaprolactone (PCL) fiber-HA hydrogel composites in contrast toblends. The second innovation is the demonstration of such ananofiber-hydrogel composite to restore soft tissue defects. Preliminarycharacterization demonstrated that the composite shared structuralcharacteristics with adipose tissue (FIG. 6) {Christman, 2012 US20120264190 A1; Young 2011. Acta Biomaterialia, 7(3), 1040-1049}. It washypothesized that this composite offers structural integrity andmechanical properties important for soft tissue regeneration. This studyhas also demonstrated the versatility and efficiency of composites, ascompared to hydrogels.

In certain aspects, a key innovation is the configuration of thenanofiber-hydrogel composite as a flat sheet that is suitable forreconstructing laminar soft tissue defects such as those posed bymissing fascia, dura, or pleura (FIG. 23).

In certain aspects, a key innovation is the composition of the hydrogelor nanofiber-hydrogel composite integrated with surgical mesh. Theresultant composition retains the low inflammatory profile and superiortissue ingrowth of the hydrogel while benefiting from the strong tensilestrength of the surgical mesh.

In certain aspects, a key innovation is the configuration of thenanofiber-hydrogel composite with linearly oriented nanofibers that canpromote preferential cellular migration (FIG. 4). This composition isideal for directing tissue regeneration along a single axis as isevident in tissues such as fascia and dura.

The successful completion of this project will deliver an off-the-shelfsolution for the restoration of missing soft tissue structures,particularly for larger defects of a laminar variety where achieving astrong, flexible sheet of material, establishing a vascular network,maintaining tissue repair site integrity, promoting cell migration andorganization, and recruiting host cells are all crucial to a sustainabletissue restoration. The extensive clinical track record for thematerials components used in this composite design, i.e. HA hydrogel andbiodegradable polyester fibers, together with these preliminary data ontissue compatibility, suggested superior tissue compatibility and astraightforward regulatory approval path for clinical translation.

Features:

In some embodiments, the invention provides the interfacial bondingbetween nanofibers and polymer network in the hydrogel component. Thisis important for the formation of a “true” composite. It wasdemonstrated that blending such fibers and hydrogel did not provide thesame degree of mechanical enhancement. There are also previous reportson the use of nanofiber-hydrogel blends. In other words, the interfacialbonding importantly differentiates this new work from the art.Furthermore, the interfacial bonding could include covalent bonds asshown in this manuscript, and secondary bonding, such as hydrogen bondsand electrostatic charge interaction.

In some embodiments, the invention provides laminar sheets of nanofibersbonded to the hydrogel in a format suitable for reconstructing laminartissues such as dura, pleura, and fascia. Optionally, the nanofibers arealigned to promote cellular ingrowth along a preferred direction.

In other embodiments, the invention provides laminar hydrogels orlaminar nanofiber-hydrogel composites reinforced by the incorporation ofstrong surgical mesh. The resultant materials can benefit from thetensile strength of the surgical mesh which can be an advantage inapplications such as abdominal wall repair while retaining the improvedinflammatory profile and superior tissue compatibility and ingrowthprovided by the hydrogels.

This is also the first work in the field that demonstrates isotropicreinforcement—the composite is stronger in all orientations, as neededto replace volumetric defects of arbitrary geometry. Designs withnanofiber mats or a small number of aligned filaments are inherentlyanisotropic. This design is capable of forming both isotropic andanisotropic materials

The work presented herein, for at least certain aspects, defines thecomponents used for the formation of composite to be a hydrogel networkwith sufficient pore size and porosity for cell migration and hosttissue ingrowth, and nanofibers which loosely include polymer fiberswith diameters ranging from 50 nm to 10 μm.

Gel/Hydrogel Component

The hydrogel composite of the invention can include any type of suitablehydrogel component. The invention contemplate nanostructure/gelcomposites that include any suitable gel component, including anysuitable hydrogel component known in the art. The gel and/or hydrogelscan be formed of any suitable synthetic or naturally-occurringmaterials.

For example, the polymer component of the gels and/or hydrogels cancomprise a cellulose ester, for example, cellulose acetate, celluloseacetate propionate (CAP), cellulose acetate butyrate (CAB), cellulosepropionate (CP), cellulose butyrate (CB), cellulose propionate butyrate(CPB), cellulose diacetate (CDA), cellulose triacetate (CTA), or thelike. These cellulose esters are described in U.S. Pat. Nos. 1,698,049,1,683,347, 1,880,808, 1,880,560, 1,984,147, 2,129,052, and 3,617,201,and may be prepared using techniques known in the art or obtainedcommercially. Commercially available cellulose esters suitable hereininclude CA 320, CA 398, CAB 381, CAB 551, CAB 553, CAP 482, CAP 504, allavailable from Eastman Chemical Company, Kingsport, Tenn. Such celluloseesters typically have a number average molecular weight of between about10,000 and about 75,000.

The cellulose esters and comprise a mixture of cellulose and celluloseester monomer units; for example, commercially available celluloseacetate butyrate contains cellulose acetate monomer units as well ascellulose butyrate monomer units and unesterified cellulose units.

The gels/hydrogels of the invention may also be comprised of otherwater-swellable polymers, such as acrylate polymers, which are generallyformed from acrylic acid, methacrylic acid, methyl acrylate, ethylacrylate, methyl methacrylate, ethyl methacrylate, and/or other vinylmonomers. Suitable acrylate polymers are those copolymers availableunder the tradename “Eudragit” from Rohm Pharma (Germany), as indicatedsupra. The Eudragit series E, L, S, RL, RS and NE copolymers areavailable as solubilized in organic solvent, in an aqueous dispersion,or as a dry powder. Preferred acrylate polymers are copolymers ofmethacrylic acid and methyl methacrylate, such as the Eudragit L andEudragit S series polymers. Particularly preferred such copolymers areEudragit L-30D-55 and Eudragit L-100-55 (the latter copolymer is aspray-dried form of Eudragit L-30D-55 that can be reconstituted withwater). The molecular weight of the Eudragit L-30D-55 and EudragitL-100-55 copolymer is approximately 135,000 Da, with a ratio of freecarboxyl groups to ester groups of approximately 1:1. The copolymer isgenerally insoluble in aqueous fluids having a pH below 5.5. Anotherparticularly suitable methacrylic acid-methyl methacrylate copolymer isEudragit S-100, which differs from Eudragit L-30D-55 in that the ratioof free carboxyl groups to ester groups is approximately 1:2. EudragitS-100 is insoluble at pH below 5.5, but unlike Eudragit L-30D-55, ispoorly soluble in aqueous fluids having a pH in the range of 5.5 to 7.0.This copolymer is soluble at pH 7.0 and above. Eudragit L-100 may alsobe used, which has a pH-dependent solubility profile between that ofEudragit L-30D-55 and Eudragit S-100, insofar as it is insoluble at a pHbelow 6.0. It will be appreciated by those skilled in the art thatEudragit L-30D-55, L-100-55, L-100, and S-100 can be replaced with otheracceptable polymers having similar pH-dependent solubilitycharacteristics.

Any of the herein-described gel/hydrogel compositions may be modified soas to contain an active agent and thereby act as an active agentdelivery system when applied to a body surface (e.g., a site of tissuerepair) in active agent-transmitting relation thereto. The release ofactive agents “loaded” into the present hydrogel compositions of theinvention typically involves both absorption of water and desorption ofthe agent via a swelling-controlled diffusion mechanism. Activeagent-containing hydrogel compositions may be employed, by way ofexample, in transdermal drug delivery systems, in wound dressings, intopical pharmaceutical formulations, in implanted drug delivery systems,in oral dosage forms, and the like.

Suitable active agents that may be incorporated into the presenthydrogel compositions and delivered systemically (e.g., with atransdermal, oral, or other dosage form suitable for systemicadministration of a drug) include, but are not limited to: analepticagents; analgesic agents; anesthetic agents; antiarthritic agents;respiratory drugs, including antiasthmatic agents; anticancer agents,including antineoplastic drugs; anticholinergics; anticonvulsants;antidepressants; antidiabetic agents; antidiarrheals; antihelminthics;antihistamines; antihyperlipidemic agents; antihypertensive agents;anti-infective agents such as antibiotics and antiviral agents;antiinflammatory agents; antimigraine preparations; antinauseants;antiparkinsonism drugs; antipruritics; antipsychotics; antipyretics;antispasmodics; antitubercular agents; antiulcer agents; antiviralagents; anxiolytics; appetite suppressants; attention deficit disorder(ADD) and attention deficit hyperactivity disorder (ADHD) drugs;cardiovascular preparations including calcium channel blockers,antianginal agents, central nervous system (CNS) agents, beta-blockersand antiarrhythmic agents; central nervous system stimulants; cough andcold preparations, including decongestants; diuretics; geneticmaterials; herbal remedies; hormonolytics; hypnotics; hypoglycemicagents; immunosuppressive agents; leukotriene inhibitors; mitoticinhibitors; muscle relaxants; narcotic antagonists; nicotine;nutritional agents, such as vitamins, essential amino acids and fattyacids; ophthalmic drugs such as antiglaucoma agents; parasympatholytics;peptide drugs; psychostimulants; sedatives; steroids, includingprogestogens, estrogens, corticosteroids, androgens and anabolic agents;smoking cessation agents; sympathomimetics; tranquilizers; andvasodilators including general coronary, peripheral and cerebral.Specific active agents with which the present adhesive compositions areuseful include, without limitation, anabasine, capsaicin, isosorbidedinitrate, aminostigmine, nitroglycerine, verapamil, propranolol,silabolin, foridone, clonidine, cytisine, phenazepam, nifedipine,fluacizin, and salbutamol.

For topical drug administration and/or medicated cushions (e.g.,medicated footpads), suitable active agents include, by way of example,the following:

Bacteriostatic and bactericidal agents: Suitable bacteriostatic andbactericidal agents include, by way of example: halogen compounds suchas iodine, iodopovidone complexes (i.e., complexes of PVP and iodine,also referred to as “povidine” and available under the tradenameBetadine from Purdue Frederick), iodide salts, chloramine,chlorohexidine, and sodium hypochlorite; silver and silver-containingcompounds such as sulfadiazine, silver protein acetyltannate, silvernitrate, silver acetate, silver lactate, silver sulfate and silverchloride; organotin compounds such as tri-n-butyltin benzoate; zinc andzinc salts; oxidants, such as hydrogen peroxide and potassiumpermanganate; aryl mercury compounds, such as phenylmercury borate ormerbromin; alkyl mercury compounds, such as thiomersal; phenols, such asthymol, o-phenyl phenol, 2-benzyl-4-chlorophenol, hexachlorophen andhexylresorcinol; and organic nitrogen compounds such as8-hydroxyquinoline, chlorquinaldol, clioquinol, ethacridine, hexetidine,chlorhexedine, and ambazone.

Antibiotic agents: Suitable antibiotic agents include, but are notlimited to, antibiotics of the lincomycin family (referring to a classof antibiotic agents originally recovered from Streptomyceslincolnensis), antibiotics of the tetracycline family (referring to aclass of antibiotic agents originally recovered from Streptomycesaureofaciens), and sulfur-based antibiotics, i.e., sulfonamides.Exemplary antibiotics of the lincomycin family include lincomycin,clindamycin, related compounds as described, for example, in U.S. Pat.Nos. 3,475,407, 3,509,127, 3,544,551 and 3,513,155, andpharmacologically acceptable salts and esters thereof. Exemplaryantibiotics of the tetracycline family include tetracycline itself,chlortetracycline, oxytetracycline, tetracycline, demeclocycline,rolitetracycline, methacycline and doxycycline and theirpharmaceutically acceptable salts and esters, particularly acid additionsalts such as the hydrochloride salt. Exemplary sulfur-based antibioticsinclude, but are not limited to, the sulfonamides sulfacetamide,sulfabenzamide, sulfadiazine, sulfadoxine, sulfamerazine,sulfamethazine, sulfamethizole, sulfamethoxazole, and pharmacologicallyacceptable salts and esters thereof, e.g., sulfacetamide sodium.

Pain relieving agents: Suitable pain relieving agents are localanesthetics, including, but not limited to, acetamidoeugenol, alfadoloneacetate, alfaxalone, amucaine, amolanone, amylocaine, benoxinate,betoxycaine, biphenamine, bupivacaine, burethamine, butacaine, butaben,butanilicaine, buthalital, butoxycaine, carticaine, 2-chloroprocaine,cinchocaine, cocaethylene, cocaine, cyclomethycaine, dibucaine,dimethisoquin, dimethocaine, diperadon, dyclonine, ecgonidine, ecgonine,ethyl aminobenzoate, ethyl chloride, etidocaine, etoxadrol,.beta.-eucaine, euprocin, fenalcomine, fomocaine, hexobarbital,hexylcaine, hydroxydione, hydroxyprocaine, hydroxytetracaine, isobutylp-aminobenzoate, kentamine, leucinocaine mesylate, levoxadrol,lidocaine, mepivacaine, meprylcaine, metabutoxycaine, methohexital,methyl chloride, midazolam, myrtecaine, naepaine, octacaine, orthocaine,oxethazaine, parethoxycaine, phenacaine, phencyclidine, phenol,piperocaine, piridocaine, polidocanol, pramoxine, prilocaine, procaine,propanidid, propanocaine, proparacaine, propipocaine, propofol,propoxycaine, pseudococaine, pyrrocaine, risocaine, salicyl alcohol,tetracaine, thialbarbital, thimylal, thiobutabarbital, thiopental,tolycaine, trimecaine, zolamine, and combinations thereof. Tetracaine,lidocaine and prilocaine are referred pain relieving agents herein.

Other topical agents that may be delivered using the present hydrogelcompositions as drug delivery systems include the following: antifungalagents such as undecylenic acid, tolnaftate, miconazole, griseofulvine,ketoconazole, ciclopirox, clotrimazole and chloroxylenol; keratolyticagents, such as salicylic acid, lactic acid and urea; vessicants such ascantharidin; anti-acne agents such as organic peroxides (e.g., benzoylperoxide), retinoids (e.g., retinoic acid, adapalene, and tazarotene),sulfonamides (e.g., sodium sulfacetamide), resorcinol, corticosteroids(e.g., triamcinolone), alpha-hydroxy acids (e.g., lactic acid andglycolic acid), alpha-keto acids (e.g., glyoxylic acid), andantibacterial agents specifically indicated for the treatment of acne,including azelaic acid, clindamycin, erythromycin, meclocycline,minocycline, nadifloxacin, cephalexin, doxycycline, and ofloxacin;skin-lightening and bleaching agents, such as hydroquinone, kojic acid,glycolic acid and other alpha-hydroxy acids, artocarpin, and certainorganic peroxides; agents for treating warts, including salicylic acid,imiquimod, dinitrochlorobenzene, dibutyl squaric acid, podophyllin,podophyllotoxin, cantharidin, trichloroacetic acid, bleomycin,cidofovir, adefovir, and analogs thereof; and anti-inflammatory agentssuch as corticosteroids and nonsteroidal anti-inflammatory drugs(NSAIDs), where the NSAIDS include ketoprofen, flurbiprofen, ibuprofen,naproxen, fenoprofen, benoxaprofen, indoprofen, pirprofen, carprofen,oxaprozin, pranoprofen, suprofen, alminoprofen, butibufen, fenbufen, andtiaprofenic acid.

For wound dressings, suitable active agents are those useful for thetreatment of wounds, and include, but are not limited to bacteriostaticand bactericidal compounds, antibiotic agents, pain relieving agents,vasodilators, tissue-healing enhancing agents, amino acids, proteins,proteolytic enzymes, cytokines, and polypeptide growth factors.

For topical and transdermal administration of some active agents, and inwound dressings, it may be necessary or desirable to incorporate apermeation enhancer into the hydrogel composition in order to enhancethe rate of penetration of the agent into or through the skin. Suitableenhancers include, for example, the following: sulfoxides such asdimethylsulfoxide (DMSO) and decylmethylsulfoxide; ethers such asdiethylene glycol monoethyl ether (available commercially as Transcutol)and diethylene glycol monomethyl ether; surfactants such as sodiumlaurate, sodium lauryl sulfate, cetyltrimethylammonium bromide,benzalkonium chloride, Poloxamer (231, 182, 184), Tween (20, 40, 60, 80)and lecithin (U.S. Pat. No. 4,783,450); the 1-substitutedazacycloheptan-2-ones, particularly 1-n-dodecylcyclaza-cycloheptan-2-one(available under the trademark Azone from Nelson Research & DevelopmentCo., Irvine, Calif.; see U.S. Pat. Nos. 3,989,816, 4,316,893, 4,405,616and 4,557,934); alcohols such as ethanol, propanol, octanol, decanol,benzyl alcohol, and the like; fatty acids such as lauric acid, oleicacid and valeric acid; fatty acid esters such as isopropyl myristate,isopropyl palmitate, methylpropionate, and ethyl oleate; polyols andesters thereof such as propylene glycol, ethylene glycol, glycerol,butanediol, polyethylene glycol, and polyethylene glycol monolaurate(PEGML; see, e.g., U.S. Pat. No. 4,568,343); amides and othernitrogenous compounds such as urea, dimethylacetamide (DMA),dimethylformamide (DMF), 2-pyrrolidone, 1-methyl-2-pyrrolidone,ethanolamine, diethanolamine and triethanolamine; terpenes; alkanones;and organic acids, particularly salicylic acid and salicylates, citricacid and succinic acid. Mixtures of two or more enhancers may also beused.

In certain other embodiments, the composite compositions of theinvention comprising a gel (e.g., a hydrogel component) and ananostructure may also comprise additional optional additive components.Such components are known in the art and can include, for example,fillers, preservatives, pH regulators, softeners, thickeners, pigments,dyes, refractive particles, stabilizers, toughening agents,detackifiers, pharmaceutical agents (e.g., antibiotics, angiogenesispromoters, antifungal agents, immunosuppressing agents, antibodies, andthe like), and permeation enhancers. These additives, and amountsthereof, are selected in such a way that they do not significantlyinterfere with the desired chemical and physical properties of thehydrogel composition.

Absorbent fillers may be advantageously incorporated to control thedegree of hydration when the adhesive is on the skin or other bodysurface. Such fillers can include microcrystalline cellulose, talc,lactose, kaolin, mannitol, colloidal silica, alumina, zinc oxide,titanium oxide, magnesium silicate, magnesium aluminum silicate,hydrophobic starch, calcium sulfate, calcium stearate, calciumphosphate, calcium phosphate dihydrate, woven and non-woven paper andcotton materials. Other suitable fillers are inert, i.e., substantiallynon-adsorbent, and include, for example, polyethylenes, polypropylenes,polyurethane polyether amide copolymers, polyesters and polyestercopolymers, nylon and rayon.

The compositions can also include one or more preservatives.Preservatives include, by way of example, p-chloro-m-cresol, phenylethylalcohol, phenoxyethyl alcohol, chlorobutanol, 4-hydroxybenzoic acidmethylester, 4-hydroxybenzoic acid propylester, benzalkonium chloride,cetylpyridinium chloride, chlorohexidine diacetate or gluconate,ethanol, and propylene glycol.

The compositions may also include pH regulating compounds. Compoundsuseful as pH regulators include, but are not limited to, glycerolbuffers, citrate buffers, borate buffers, phosphate buffers, or citricacid-phosphate buffers may also be included so as to ensure that the pHof the hydrogel composition is compatible with that of an individual'sbody surface.

The compositions may also include suitable softening agents. Suitablesofteners include citric acid esters, such as triethylcitrate or acetyltriethylcitrate, tartaric acid esters such as dibutyltartrate, glycerolesters such as glycerol diacetate and glycerol triacetate; phthalic acidesters, such as dibutyl phthalate and diethyl phthalate; and/orhydrophilic surfactants, preferably hydrophilic non-ionic surfactants,such as, for example, partial fatty acid esters of sugars, polyethyleneglycol fatty acid esters, polyethylene glycol fatty alcohol ethers, andpolyethylene glycol sorbitan-fatty acid esters.

The compositions may also include thickening agents. Preferredthickeners herein are naturally occurring compounds or derivativesthereof, and include, by way of example: collagen; galactomannans;starches; starch derivatives and hydrolysates; cellulose derivativessuch as methyl cellulose, hydroxypropylcellulose, hydroxyethylcellulose, and hydroxypropyl methyl cellulose; colloidal silicic acids;and sugars such as lactose, saccharose, fructose and glucose. Syntheticthickeners such as polyvinyl alcohol,vinylpyrrolidone-vinylacetate-copolymers, polyethylene glycols, andpolypropylene glycols may also be used.

In certain embodiments, the hydrogel composite of the inventioncomprising a hydrogel and a nanostructure further comprises a componentthat promotes angiogenesis. A challenge to achieving clinically relevantsoft tissue regeneration prior to the present invention is that theregenerated tissue preferably should be re-vascularized. Therefore, anymaterial that promotes soft tissue regeneration preferably should alsoencourage angiogenesis. One way to achieve this is through the use ofheparin-containing hydrogel components, which can serve as growth factorbinding sites to enrich and retain growth factors promoting angiogenesisand tissue formation.

In various other embodiments, the composite materials of the inventioncan be based on hyaluronic acid (HA) as they hydrogel material. HA is anon-sulfated, linear polysaccharide with repeating disaccharide unitswhich form the hydrogel component. HA is also a non-immunogenic, nativecomponent of the extracellular matrix in human tissues, and widely usedas a dermal filler in aesthetic and reconstructive procedures.

Breakdown of HA is facilitated by native hyaluronidases whose expressionis increased in areas of tissue damage and inflammation. Importantly,studies have shown that small HA degradation fragments of 3-10disaccharide units are potent regulators of endothelial cellproliferation, migration, tubule formation, and angiogenesis. Thesebiological functions of HA are thought to be mediated via CD44 in apathway involving Ras and PKC. Blockade of CD44/HA interactions usinganti-CD44 antibodies reduced proliferation and migration of humanmicrovascular endothelial cells in vitro. HA hydrogels have beeninvestigated as potential matrices for cell delivery in a variety ofmodels of cell and tissue injury. These hydrogels can serve as aprotective and supporting scaffold for cells and can also reducescarring. Thus, it is believed HA has a critical role in enhancingtissue regeneration by promoting cell infiltration and promotingangiogenesis.

First, the material has three-dimensional integrity and a consistencysimilar to that of native fat tissue. This renders it suitable foroff-the-shelf restoration of missing soft tissue volume. Second, thematerial preferably may be deposited with a plurality of flexiblenanofibers that can serve as substrates for migration of adipocytes andendothelial progenitors. Third, the material has sufficient porosity toallow these precursor cells to rapidly infiltrate and integrate into thescaffold rather than forming a fibrous capsule around it. Fourth, the HAhydrogel component provides compressibility and volumetric expansionwhile also providing important angiogenic cues. Fifth, the nanofiber andhydrogel components are biodegradable allowing them to be replaced byregenerated soft tissue. Sixth, all component materials have strongsafety track records in numerous FDA-approved devices, potentiallyreducing regulatory hurdles for clinical translation.

The gel/hydrogel/nanostructure composites of the invention can alsoinclude tissue-repairing agents, such as, a number of growth factors,including epidermal growth factor (EDF), PDGF, and nerve growth factors(NGF's). For example, the compositions may include EGF. Epidermal GrowthFactor (EGF) was discovered after the observation that cutaneous woundsin laboratory mice seemed to heal more rapidly when the mice wereallowed to lick them. This was not simply due to some antiseptic agentin saliva (such as lysozyme). A specific growth factor, now known asEGF, was shown to be responsible. EGF is identical to urogastrone, andhas angiogenic properties. Transforming growth factor-alpha(TGF-.alpha.) is very similar, binding to the same receptor and is evenmore effective in stimulating epithelial cell regeneration(epithelisation).

Thus, hydrogels of the present invention comprising EGF/TGF mayadvantageously be used in the acceleration of wound healing and burns,reduction in keloid scar formation (especially for burns), skinengraftment dressings, and the treatment of chronic leg ulcers.

Tissue-repairing agents useful in the present invention include a numberof growth factors, including epidermal growth factor (EDF), PDGF, andnerve growth factors (NGF's). Generally, growth-promoting hormones willaffect between one and four tissues. Many of the products developed fromsuch proteins are targeted towards wound repairs of one kind or another,although there are other indications. Some of the most important tissuegrowth factors are described further below.

The gel/nanostructure compositions of the invention may also include oneor more growth factors that may be useful in the tissue repair methodsand other applications of the invention.

For example, the invention contemplates include PDGF in the compositionsof the invention. Platelet-Derived Growth Factor (PDGF) is a mitogen foralmost all mesenchymally-derived cells, i.e. blood, muscle, bone,cartilage, and connective tissue cells. It is a dimeric glycoproteinexisting as AA or BB homodimers, or as the AB heterodimer. As with manygrowth factors, PDGF is now considered to be a member of a larger familyof factors. In addition to PDGF, this family includes the homodimericfactors vascular endothelial growth factor (VEGF) and placental growthfactor (PIGF), VEGF/PIGF heterodimers, and connective tissue growthfactor (CTGF), a PDGF-like factor secreted by human vascular endothelialcells and fibroblasts. Along with NGF, TGF-.beta. and glycoproteinhormones such as human chorionic gonadotropic hormone (hCG), PDGF is nowclassified as a member of the cysteine-knot growth factor superfamily.All of these factors may be used in conjunction with hydrogels of thepresent invention.

PDGF is produced by platelets and released in the course of bloodclotting. It is just one of the growth factors that derive from thesecells. PDGF attracts fibroblasts and white blood cells to the site ofthe injury, as well as stimulating the growth of replacement connectivetissue (mainly fibroblasts and smooth muscle cells). It stimulates celldivision in various cells, including those that produce collagen, soencouraging angiogenesis. It also stimulates mitogenesis,vasoconstriction, chemotaxis, enzyme activity and calcium mobilization.

Blood platelet derived growth factors may be used to restore bone andsoft tissue regrowth during certain treatments using the compositions ofthe invention and to accelerate the healing process of chronic and acutewounds. Accordingly, hydrogel/nanostructure compositions of the presentinvention may advantageously comprise a platelet derived growth factorcocktail.

Hydrogel/nanostructure compositions of the present invention may be usedin gene therapy for local delivery of the PDGF gene, for example.Plasmid DNA encoding PDGF is incorporated into the hydrogel matrix andgranulation tissue fibroblasts, which originate in viable tissuesurrounding the wound, proliferate and migrate into the matrix, actingas targets for plasmid gene transfer and expression.

The hydrogel/nanostructure compositions of the invention may alsoinclude VEGF to promote angiogenesis. Vascular Endothelial Growth Factor(VEGF—also known as vascular permeability factor) is another vasculargrowth factor, and is a multifunctional angiogenic cytokine. Itcontributes to angiogenesis (blood vessel growth) both indirectly anddirectly by stimulating proliferation of endothelial cells at themicrovessel level, causing them to migrate and to alter their genericexpression. VEGF also makes theses endothelial cells hyperpermeable,causing them to release plasma proteins outside the vascular space,which causes changes in the area, contributing to angiogenesis.

The compositions of the invention may also include FGF. FibroblastGrowth Factor (FGF) is actually a family of at least 19 14 18 kDpeptides belonging to the heparin-binding growth factors family, and aremitogenic for cultured fibroblasts and vascular endothelial cells. Theyare also angiogenic in vivo and this angiogenicity is enhanced by TNF.FGF's may be used in a similar manner to EGF. bFGF, also known as FGF-2,is involved in controlling human megakaryocytopoiesis and FGFs have beenshown to be effective in stimulating endothelial cell formation, and inassisting in connective tissue repair.

Hydrogel/nanostructure compositions may also comprise KeratinocyteGrowth Factor (KGF), also known as FGF-7, for use in wound healing andother disorders involving epithelial cell destruction.

Transforming Growth Factors (TGFs) have the ability to transform variouscell lines, and can confer, for example, the ability to grow in culturefor more than a limited number of generations, growth in multiple layersrather than monolayers, and the acquisition of an abnormal karyotype.There are at least five members of the TGF family, the two most widelystudied being TGF-alpha and TGF-beta. The former is mitogenic forfibroblasts and endothelial cells, angiogenic, and promotes boneresorption. Compositions also may include TGF. TGF-beta is a generalmediator of cell regulation, a powerful inhibitor of cell growth, andinhibits the proliferation of many cell types. TGF-beta can antagonisethe mitogenic effects of other peptide growth factors, and can alsoinhibit the growth of many tumour cell lines. TGF-beta also hasangiogenic effects, and promotes collagen formation in fibroblasts.Indications for hydrogels of the present invention include chronic skinulcers, such as neurotrophic foot ulcers in diabetic patients. Otherareas include wound healing, bone repair and immunosuppressive diseases.

Hydrogel/nanostructure compositions of the present invention may be usedto carry suitable cells, for example. These may be incorporated into thegel just prior to application to a wound, or other suitable area, tomaximise efficacy. Suitable cells include autologous fibroblasts andkeratinocytes, which are mainly responsible for dermis and epidermisformation. Separate gels each comprising one cell type may be appliedconsecutively or together, or one gel may comprise both cell types, butthis is generally less preferred.

Hydrogel/nanostructure compositions of the present invention mayusefully comprise collagen, for example. Although collagen, in thisform, is unlikely to serve a useful structural function, it primarilyserves as a sacrificial protein where proteolytic activity isundesirably high, thereby helping to prevent maceration of healthytissue, for example.

Hydrogel/nanostructure compositions can also include certain enzymes.Enzymes are used in the debridement of both acute and chronic wounds.Debridement is the removal of nonviable tissue and foreign matter from awound and is a naturally occurring event in the wound-repair process.During the inflammatory phase, neutrophils and macrophages digest andremove “used” platelets, cellular debris, and avascular injured tissuefrom the wound area. However, with the accumulation of significantamounts of damaged tissue, this natural process becomes overwhelmed andinsufficient. Build-up of necrotic tissue then places considerablephagocytic demand on the wound and retards wound healing. Consequently,debridement of necrotic tissue is a particular objective of topicaltherapy and an important component of optimal wound management.

Enzymes, for example, may be incorporated into hydrogels of the presentinvention for topical application to provide a selective method ofdebridement. Suitable enzymes may be derived from various sources, suchas krill, crab, papaya, bovine extract, and bacteria Commerciallyavailable, suitable enzymes include collagenase, papain/urea, and afibrinolysin and deoxyribonuclease combination.

Enzymes for use in the present invention generally work in one of twoways: by directly digesting the components of slough (e.g., fibrin,bacteria, leukocytes, cell debris, serous exudate, DNA); or, bydissolving the collagen “anchors” that secure the avascular tissue tothe underlying wound bed.

Hydrogels of the present invention may comprise Dakin's solution, ifdesired, generally to exert antimicrobial effects and odour control. Asa debridement agent, Dakin's solution is non-selective because of itscytotoxic properties. Dakin's solution denatures protein, rendering itmore easily removed from the wound. Loosening of the slough alsofacilitates debridement by other methods. Hydrogels comprising Dakin'ssolution may be changed twice daily if the goal is debridement.Periwound skin protection should generally be provided with ointments,liquid skin barrier film dressings, or solid skin barrier wafers, forexample.

The gel of the present invention may be delivered by any suitablemethod, such as via a syringe or bellows pack (single dose deliverysystems) or a multidose system, such as a pressurised delivery system ordelivery via a ‘bag in the can’ type system (such as that published inWO98/32675). An example of a bellows pack is shown in published UKdesign number 2082665.

As such, the present invention also extends to a single dose deliverysystem comprising a gel according to the present invention, for thetreatment of wounds. The invention also extends to a pressuriseddelivery system comprising a gel according to the present invention, anda pressurised hydrogel according to the present invention in an aerosolcontainer capable of forming a spray upon release of pressure therefrom.Use of such delivery means allows the gel to be delivered to areas on apatient which are otherwise difficult to reach by direct application,such as on the back of a patient when the patient is lying down.

In certain embodiment, it may be advantageous to render the hydrogelcompositions of the invention electrically conductive for use inbiomedical electrodes and other electrotherapy contexts, i.e., to attachan electrode or other electrically conductive member to the bodysurface. For example, the hydrogel composition may be used to attach atranscutaneous nerve stimulation electrode, an electrosurgical returnelectrode, or an EKG electrode to a patient's skin or mucosal tissue.These applications involve modification of the hydrogel composition soas to contain a conductive species. Suitable conductive species areionically conductive electrolytes, particularly those that are normallyused in the manufacture of conductive adhesives used for application tothe skin or other body surface, and include ionizable inorganic salts,organic compounds, or combinations of both. Examples of ionicallyconductive electrolytes include, but are not limited to, ammoniumsulfate, ammonium acetate, monoethanolamine acetate, diethanolamineacetate, sodium lactate, sodium citrate, magnesium acetate, magnesiumsulfate, sodium acetate, calcium chloride, magnesium chloride, calciumsulfate, lithium chloride, lithium perchlorate, sodium citrate andpotassium chloride, and redox couples such as a mixture of ferric andferrous salts such as sulfates and gluconates. Preferred salts arepotassium chloride, sodium chloride, magnesium sulfate, and magnesiumacetate, and potassium chloride is most preferred for EKG applications.Although virtually any amount of electrolyte may be present in theadhesive compositions of the invention, it is preferable that anyelectrolyte present be at a concentration in the range of about 0.1 toabout 15 wt. % of the hydrogel composition. The procedure described inU.S. Pat. No. 5,846,558 to Nielsen et al. for fabricating biomedicalelectrodes may be adapted for use with the hydrogel compositions of theinvention, and the disclosure of that patent is incorporated byreference with respect to manufacturing details. Other suitablefabrication procedures may be used as well, as will be appreciated bythose skilled in the art.

Crosslinking

For certain applications, particularly when high cohesive strength isdesired, the polymers of the gel/hydrogels of the invention may becovalently crosslinked. The disclosure contemplates that crosslinkingmay be desired as between the polymers of the gel/hydrogel component,but also crosslinking may be desired as between the polymers of thegel/hydrogel and the nanostructure components of the composite materialsof the invention. The invention contemplates any suitable means forcrosslinking polymers to one another, and crosslinking the gel/hydrogelpolymers with the nanostructure components of the invention. Thegel/hydrogel polymers may be covalently crosslinked to other polymers orto the nanostructures, either intramolecularly or intermolecularly orthrough covalent bonds. In the former case, there are no covalent bondslinking the polymers to one another or to the nanostructures, while inthe latter case, there are covalent crosslinks binding the polymers toone another or to the nanostructures. The crosslinks may be formed usingany suitable means, including using heat, radiation, or a chemicalcuring (crosslinking) agent. The degree of crosslinking should besufficient to eliminate or at least minimize cold flow undercompression. Crosslinking also includes the use of a third molecule, a“cross-linker” utilized in the cross-linking process.

For thermal crosslinking, a free radical polymerization initiator isused, and can be any of the known free radical-generating initiatorsconventionally used in vinyl polymerization. Preferred initiators areorganic peroxides and azo compounds, generally used in an amount fromabout 0.01 wt. % to 15 wt. %, preferably 0.05 wt. % to 10 wt. %, morepreferably from about 0.1 wt. % to about 5% and most preferably fromabout 0.5 wt. % to about 4 wt. % of the polymerizable material. Suitableorganic peroxides include dialkyl peroxides such as t-butyl peroxide and2,2bis(t-butylperoxy)propane, diacyl peroxides such as benzoyl peroxideand acetyl peroxide, peresters such as t-butyl perbenzoate and t-butylper-2-ethylhexanoate, perdicarbonates such as dicetyl peroxy dicarbonateand dicyclohexyl peroxy dicarbonate, ketone peroxides such ascyclohexanone peroxide and methylethylketone peroxide, andhydroperoxides such as cumene hydroperoxide and tert-butylhydroperoxide. Suitable azo compounds include azo bis (isobutyronitrile)and azo bis (2,4-dimethylvaleronitrile). The temperature for thermallycrosslinking will depend on the actual components and may be readilydeduced by one of ordinary skill in the art, but typically ranges fromabout 80 C. to about 200 C.

Crosslinking may also be accomplished with radiation, typically in thepresence of a photoinitiator. The radiation may be ultraviolet, alpha,beta, gamma, electron beam, and x-ray radiation, although ultravioletradiation is preferred. Useful photosensitizers are triplet sensitizersof the “hydrogen abstraction” type, and include benzophenone andsubstituted benzophenone and acetophenones such as benzyl dimethylketal, 4-acryloxybenzophenone (ABP), 1-hydroxy-cyclohexyl phenyl ketone,2,2-diethoxyacetophenone and 2,2-dimethoxy-2-phenylaceto-phenone,substituted alpha-ketols such as 2-methyl-2-hydroxypropiophenone,benzoin ethers such as benzoin methyl ether and benzoin isopropyl ether,substituted benzoin ethers such as anisoin methyl ether, aromaticsulfonyl chlorides such as 2-naphthalene sulfonyl chloride, photoactiveoximes such as 1-phenyl-1,2-propanedione-2-(O-ethoxy-carbonyl)-oxime,thioxanthones including alkyl- and halogen-substituted thioxanthonsesuch as 2-isopropylthioxanthone, 2-chlorothioxanthone, 2,4 dimethylthioxanone, 2,4 dichlorothioxanone, and 2,4-diethyl thioxanone, and acylphosphine oxides. Radiation having a wavelength of 200 to 800 nm,preferably, 200 to 500 nm, is preferred for use herein, and lowintensity ultraviolet light is sufficient to induce crosslinking in mostcases. However, with photosensitizers of the hydrogen abstraction type,higher intensity UV exposure may be necessary to achieve sufficientcrosslinking. Such exposure can be provided by a mercury lamp processorsuch as those available from PPG, Fusion, Xenon, and others.Crosslinking may also be induced by irradiating with gamma radiation oran electron beam. Appropriate irradiation parameters, i.e., the type anddose of radiation used to effect crosslinking, will be apparent to thoseskilled in the art.

Suitable chemical curing agents, also referred to as chemicalcross-linking “promoters,” include, without limitation, polymercaptanssuch as 2,2-dimercapto diethylether, dipentaerythritolhexa(3-mercaptopropionate), ethylene bis(3-mercaptoacetate),pentaerythritol tetra(3-mercaptopropionate), pentaerythritoltetrathioglycolate, polyethylene glycol dimercaptoacetate, polyethyleneglycol di(3-mercaptopropionate), trimethylolethanetri(3-mercaptopropionate), trimethylolethane trithioglycolate,trimethylolpropane tri(3-mercaptopropionate), trimethylolpropanetrithioglycolate, dithioethane, di- or trithiopropane and 1,6-hexanedithiol. The crosslinking promoter is added to the uncrosslinkedhydrophilic polymer to promote covalent crosslinking thereof, or to ablend of the uncrosslinked hydrophilic polymer and the complementaryoligomer, to provide crosslinking between the two components.

The polymers and/or nanostructures may also be crosslinked prior toadmixture with the complementary oligomer. In such a case, it may bepreferred to synthesize the polymer in crosslinked form, by admixing amonomeric precursor to the polymer with multifunctional comonomer andcopolymerizing. Examples of monomeric precursors and correspondingpolymeric products are as follows: N-vinyl amide precursors for apoly(N-vinyl amide) product; N-alkylacrylamides for apoly(N-alkylacrylamide) product; acrylic acid for a polyacrylic acidproduct; methacrylic acid for a polymethacrylic acid product;acrylonitrile for a poly(acrylonitrile) product; and N-vinyl pyrrolidone(NVP) for a poly(vinylpyrrolidone) (PVP) product. Polymerization may becarried out in bulk, in suspension, in solution, or in an emulsion.Solution polymerization is preferred, and polar organic solvents such asethyl acetate and lower alkanols (e.g., ethanol, isopropyl alcohol,etc.) are particularly preferred. For preparation of hydrophilic vinylpolymers, synthesis will typically take place via a free radicalpolymerization process in the presence of a free radical initiator asdescribed above. The multifunctional comonomer include, for example,bisacrylamide, acrylic or methacrylic esters of diols such as butanedioland hexanediol (1,6-hexane diol diacrylate is preferred), otheracrylates such as pentaerythritol tetraacrylate, and 1,2-ethylene glycoldiacrylate, and 1,12-dodecanediol diacrylate. Other usefulmultifunctional crosslinking monomers include oligomeric and polymericmultifunctional (meth)acrylates, e.g., poly(ethylene oxide) diacrylateor poly(ethylene oxide) dimethacrylate; polyvinylic crosslinking agentssuch as substituted and unsubstituted divinylbenzene; and difunctionalurethane acrylates such as EBECRYL 270 and EBECRYL 230 (1500 weightaverage molecular weight and 5000 weight average molecular weightacrylated urethanes, respectively—both available from UCB of Smyrna,Ga.), and combinations thereof. If a chemical crosslinking agent isemployed, the amount used will preferably be such that the weight ratioof crosslinking agent to hydrophilic polymer is in the range of about1:100 to 1:5. To achieve a higher crosslink density, if desired,chemical crosslinking is combined with radiation curing.

Nanostructures

The nanostructure components of the invention may be in any suitableform including fibers, filaments, mesh sections, branched filaments ornetworks, sheets, or shaped particles. The nanostructures may alsocomprise any suitable chemical functional groups to facilitate thecovalent or noncovalent crosslinking between the nanostructures and thepolymers of the hydrogels of the invention. Method, techniques, andmaterials are well known in the art for making and functionalizingnanostructures.

In certain embodiments, microfabrication methods are used to make thenanostructures of the invention. In various embodiments, the discloseddevices can be assembled and/or manufactured using any suitablemicrofabrication technique. Such methods and techniques are widely knownin the art.

Microfabrication processes that can be used in making the nanostructuresdisclosed herein include lithography; etching techniques, such aslasers, plasma etching, photolithography, or chemical etching such aswet chemical, dry, and photoresist removal; or by solid free formtechniques, including three-dimensional printing (3DP),stereolithography (SLA), selective laser sintering (SLS), ballisticparticle manufacturing (BPM) and fusion deposition modeling (FDM); bymicromachining; thermal oxidation of silicon; electroplating andelectroless plating; diffusion processes, such as boron, phosphorus,arsenic, and antimony diffusion; ion implantation; film deposition, suchas evaporation (filament, electron beam, flash, and shadowing and stepcoverage), sputtering, chemical vapor deposition (CVD), epitaxy (vaporphase, liquid phase, and molecular beam), electroplating, screenprinting, lamination or by combinations thereof. See Jaeger,Introduction to Microelectronic Fabrication (Addison-Wesley PublishingCo., Reading Mass. 1988); Runyan, et al., Semiconductor IntegratedCircuit Processing Technology (Addison-Wesley Publishing Co., ReadingMass. 1990); Proceedings of the IEEE Micro Electro Mechanical SystemsConference 1987-1998; Rai-Choudhury, ed., Handbook of Microlithography,Micromachining & Microfabrication (SPIE Optical Engineering Press,Bellingham, Wash. 1997). The selection of the material that is used asthe mold determines how the surface is configured to form the branchingstructure.

For example, state of the art processes for fabrication of Micro ElectroMechanical Systems (MEMS) utilizing photolithographic processes andmethods derived from the semiconductor industry may be used. Morerecently developed methods include “soft lithography” (Whitesides et al,Angew chem. Int ed, 37; 550-575, (1998)) and microfluidic tectonics(U.S. Pat. No. 6,488,872, Beebe et al., Nature; 404:588-59 (2000)).Reviews and other discussions of polymer microdevice fabrication includeMadou, M. J. Fundamentals of Microfabrication: The Science ofMiniaturization; 2nd ed.; CRC Press: Boca Raton, 1997; Becker, H., andLocascio, L. E. “Polymer microfluidic devices.” Talanta, 56(2):267-287,2002; Quake, S. R., and Scherer, A. “From micro- to nanofabrication withsoft materials.” Science, 290(5496):1536-1540, 2000; and Whitesides, G.M., and Stroock, A. D. “Flexible methods for microfluidics.” PhysicsToday, 54(6):42-48, 2001, each of which are incorporated herein byreference.

The nanostructures of the invention may also be fabricated byelectrostatic spinning (also referred to as electrospinning). Thetechnique of electrospinning of liquids and/or solutions capable offorming fibers, is well known and has been described in a number ofpatents, such as, for example, U.S. Pat. Nos. 4,043,331 and 5,522,879.The process of electrospinning generally involves the introduction of aliquid into an electric field, so that the liquid is caused to producefibers. These fibers are generally drawn to a conductor at an attractiveelectrical potential for collection. During the conversion of the liquidinto fibers, the fibers harden and/or dry. This hardening and/or dryingmay be caused by cooling of the liquid, i.e., where the liquid isnormally a solid at room temperature; by evaporation of a solvent, e.g.,by dehydration (physically induced hardening); or by a curing mechanism(chemically induced hardening).

The process of electrostatic spinning has typically been directed towardthe use of the fibers to create a mat or other non-woven material, asdisclosed, for example, in U.S. Pat. No. 4,043,331. Nanofibers rangingfrom 50 nm to 5 micrometers in diameter can be electrospun into anonwoven or an aligned nanofiber mesh. Due to the small fiber diameters,electrospun textiles inherently possess a very high surface area and asmall pore size. These properties make electrospun fabrics potentialcandidates for a number of applications including: membranes, tissuescaffolding, and other biomedical applications.

Electrostatically spun fibers can be produced having very thindiameters. Parameters that influence the diameter, consistency, anduniformity of the electrospun fibers include the polymeric material andcross-linker concentration (loading) in the fiber-forming combination,the applied voltage, and needle collector distance. According to oneembodiment of the present invention, a nanofiber has a diameter rangingfrom about 1 nm to about 100.mu.m. In other embodiments, the nanofiberhas a diameter in a range of about 1 nm to about 1000 nm. Further, thenanofiber may have an aspect ratio in a range of at least about 10 toabout at least 100. It will be appreciated that, because of the verysmall diameter of the fibers, the fibers have a high surface area perunit of mass. This high surface area to mass ratio permits fiber-formingsolutions or liquids to be transformed from liquid or solvatedfiber-forming materials to solid nanofibers in fractions of a second.

The polymeric material used to form the nanofibers/nanostructures of theinvention may be selected from any fiber forming material which iscompatible with the cross-linking agents. Depending upon the intendedapplication, the fiber-forming polymeric material may be hydrophilic,hydrophobic or amphiphilic. Additionally, the fiber-forming polymericmaterial may be a thermally responsive polymeric material.

Synthetic or natural, biodegradable or non-biodegradable polymers mayform the nanofibers/nanostructures of the invention. A “syntheticpolymer” refers to a polymer that is synthetically prepared and thatincludes non-naturally occurring monomeric units. For example, asynthetic polymer can include non-natural monomeric units such asacrylate or acrylamide units. Synthetic polymers are typically formed bytraditional polymerization reactions, such as addition, condensation, orfree-radical polymerizations. Synthetic polymers can also include thosehaving natural monomeric units, such as naturally-occurring peptide,nucleotide, and saccharide monomeric units in combination withnon-natural monomeric units (for example synthetic peptide, nucleotide,and saccharide derivatives). These types of synthetic polymers can beproduced by standard synthetic techniques, such as by solid phasesynthesis, or recombinantly, when allowed.

A “natural polymer” refers to a polymer that is either naturally,recombinantly, or synthetically prepared and that consists of naturallyoccurring monomeric units in the polymeric backbone. In some cases, thenatural polymer may be modified, processed, derivatized, or otherwisetreated to change the chemical and/or physical properties of the naturalpolymer. In these instances, the term “natural polymer” will be modifiedto reflect the change to the natural polymer (for example, a“derivatized natural polymer”, or a “deglycosylated natural polymer”).

Nanofiber materials, for example, may include both addition polymer andcondensation polymer materials such as polyolefin, polyacetal,polyamide, polyester, cellulose ether and ester, polyalkylene sulfide,polyarylene oxide, polysulfone, modified polysulfone polymers andmixtures thereof. Exemplary materials within these generic classesinclude polyethylene, poly(.epsilon.-caprolactone), poly(lactate),poly(glycolate), polypropylene, poly(vinylchloride),polymethylmethacrylate (and other acrylic resins), polystyrene, andcopolymers thereof (including ABA type block copolymers),poly(vinylidene fluoride), poly(vinylidene chloride), polyvinyl alcoholin various degrees of hydrolysis (87% to 99.5%) in crosslinked andnon-crosslinked forms. Exemplary addition polymers tend to be glassy (aTg greater than room temperature). This is the case forpolyvinylchloride and polymethylmethacrylate, polystyrene polymercompositions, or alloys or low in crystallinity for polyvinylidenefluoride and polyvinyl alcohol materials.

In some embodiments of the invention the nanofiber/nanostructurematerials are polyamide condensation polymers. In more specificembodiments, the polyamide condensation polymer is a nylon polymer. Theterm “nylon” is a generic name for all long chain synthetic polyamides.Another nylon can be made by the polycondensation of epsilon caprolactamin the presence of a small amount of water. This reaction forms anylon-6 (made from a cyclic lactam—also known as epsilon-aminocaproicacid) that is a linear polyamide. Further, nylon copolymers are alsocontemplated. Copolymers can be made by combining various diaminecompounds, various diacid compounds and various cyclic lactam structuresin a reaction mixture and then forming the nylon with randomlypositioned monomeric materials in a polyamide structure. For example, anylon 6,6-6,10 material is a nylon manufactured from hexamethylenediamine and a C6 and a C10 blend of diacids. A nylon 6-6,6-6,10 is anylon manufactured by copolymerization of epsilon aminocaproic acid,hexamethylene diamine and a blend of a C6 and a C10 diacid material.

Block copolymers can also be used as nanofiber materials. In preparing acomposition for the preparation of nanofibers, a solvent system can bechosen such that both blocks are soluble in the solvent. One example isan ABA (styrene-EP-styrene) or AB (styrene-EP) polymer in methylenechloride solvent. Examples of such block copolymers are a Kraton-type ofAB and ABA block polymers including styrene/butadiene andstyrene/hydrogenated butadiene(ethylene propylene), a Pebax-type ofepsilon-caprolactam/ethylene oxide and a Sympatex-type ofpolyester/ethylene oxide and polyurethanes of ethylene oxide andisocyanates.

Addition polymers such as polyvinylidene fluoride, syndiotacticpolystyrene, copolymers of vinylidene fluoride and hexafluoropropylene,polyvinyl alcohol, polyvinyl acetate, amorphous addition polymers, suchas poly(acrylonitrile) and its copolymers with acrylic acid andmethacrylates, polystyrene, poly(vinyl chloride) and its variouscopolymers, poly(methyl methacrylate) and its various copolymers, can besolution spun with relative ease because they are soluble at lowpressures and temperatures. Highly crystalline polymer like polyethyleneand polypropylene generally require higher temperature and high pressuresolvents if they are to be solution spun.

Nanofibers can also be formed from polymeric compositions comprising twoor more polymeric materials in polymer admixture, alloy format, or in acrosslinked chemically bonded structure. Two related polymer materialscan be blended to provide the nanofiber with beneficial properties. Forexample, a high molecular weight polyvinylchloride can be blended with alow molecular weight polyvinylchloride. Similarly, a high molecularweight nylon material can be blended with a low molecular weight nylonmaterial. Further, differing species of a general polymeric genus can beblended. For example, a high molecular weight styrene material can beblended with a low molecular weight, high impact polystyrene. A Nylon-6material can be blended with a nylon copolymer such as a Nylon-6; 6,6;6,10 copolymer. Further, a polyvinyl alcohol having a low degree ofhydrolysis such as a 87% hydrolyzed polyvinyl alcohol can be blendedwith a fully or super hydrolyzed polyvinyl alcohol having a degree ofhydrolysis between 98 and 99.9% and higher. All of these materials inadmixture can be crosslinked using appropriate crosslinking mechanisms.Nylons can be crosslinked using crosslinking agents that are reactivewith the nitrogen atom in the amide linkage. Polyvinyl alcohol materialscan be crosslinked using hydroxyl reactive materials such asmonoaldehydes, such as formaldehyde, ureas, melamine-formaldehyde resinand its analogues, boric acids, and other inorganic compounds,dialdehydes, diacids, urethanes, epoxies, and other known crosslinkingagents. Crosslinking reagent reacts and forms covalent bonds betweenpolymer chains to substantially improve molecular weight, chemicalresistance, overall strength and resistance to mechanical degradation.

Biodegradable polymers can also be used in the preparation of thenanostructures of the invention. Examples of classes of syntheticpolymers that have been studied as biodegradable materials includepolyesters, polyamides, polyurethanes, polyorthoesters, polycaprolactone(PCL), polyiminocarbonates, aliphatic carbonates, polyphosphazenes,polyanhydrides, and copolymers thereof. Specific examples ofbiodegradable materials that can be used in connection with, forexample, implantable medical devices include polylactide, polyglycolide,polydioxanone, poly(lactide-co-glycolide),poly(glycolide-co-polydioxanone), polyanhydrides,poly(glycolide-co-trimethylene carbonate), andpoly(glycolide-co-caprolactone). Blends of these polymers with otherbiodegradable polymers can also be used.

In some embodiments, the nanofibers are non-biodegradable polymers.Non-biodegradable refers to polymers that are generally not able to benon-enzymatically, hydrolytically or enzymatically degraded. Forexample, the non-biodegradable polymer is resistant to degradation thatmay be caused by proteases. Non-biodegradable polymers may includeeither natural or synthetic polymers.

The inclusion of cross-linking agents within the composition forming thenanofiber, allows the nanofiber to be compatible with a wide range ofsupport surfaces. The cross-linking agents can be used alone or incombination with other materials to provide a desired surfacecharacteristic.

Suitable cross-linking agents include either monomeric (small moleculematerials) or polymeric materials having at least two latent reactiveactivatable groups that are capable of forming covalent bonds with othermaterials when subjected to a source of energy such as radiation,electrical or thermal energy. In general, latent reactive activatablegroups are chemical entities that respond to specific applied externalenergy or stimuli to generate active species with resultant covalentbonding to an adjacent chemical structure. Latent reactive groups arethose groups that retain their covalent bonds under storage conditionsbut that form covalent bonds with other molecules upon activation by anexternal energy source. In some embodiments, latent reactive groups formactive species such as free radicals. These free radicals may includenitrenes, carbine or excited states of ketones upon absorption ofexternally applied electric, electrochemical or thermal energy. Variousexamples of known or commercially available latent reactive groups arereported in U.S. Pat. Nos. 4,973,493; 5,258,041; 5,563,056; 5,637,460;or 6,278,018.

For example, the commercially available multifunctionalphotocrosslinkers based on trichloromethyl triazine available eitherfrom Aldrich Chemicals, Produits Chimiques Auxiliaires et de Syntheses,(Longjumeau, France), Shin-Nakamara Chemical, Midori Chemicals Co., Ltd.or Panchim S. A. (France) can be used. The eight compounds include2,4,6-tris(trichloromethyl)-1,3,5 triazine,2-(methyl)-4,6-bis(trichloromethyl)-1,3,5-triazine,2-(4-methoxynaphthyl)-4,6-bis(trichloromethyl)-1,3,5-triazine,2-(4-ethoxynaphthyl)-4,6-bis(trichloromethyl)-1,3,5-triazine,4-(4-carboxylphenyl)-2,6-bis(trichloromethyl)-1,3,5-triazine,2-(4-methoxyphenyl)-4,6-bis(trichloromethyl)-1,3,5-triazine,2-(1-ethen-2-2′-furyl)-4,6-bis(trichloromethyl)-1,3,5-triazine and2-(4-methoxystyryl)-4,6-bis(trichloromethyl)-1,3,5-triazine.

Methods of Use and Exemplary Embodiments

The gel/hydrogel/nanostructure compositions of the invention can be usedadvantageously in numerous tissue repair situations, as well as in otherapplications, such as providing coatings on catheters and other surgicaldevices and implants. The gel/hydrogel/nanostructure compositions of theinvention can also be used to deliver active agents described herein,such as antibiotics, growth factors, and immunosuppressive agents.

In certain embodiments, the invention provides a method for healing asoft tissue defect comprising applying a composite material to a softtissue defect, wherein the composite material includes a gel and ananostructure disposed within the gel.

It will be appreciated that advantageous properties of thehydrogels/nanostructure compositions described herein include theability to: 1) provide easy characterization and quality control; 2)integrate with existing tissue matrices; 3) directly incorporate intonewly formed matrices; 4) directly include cells and bioactive factors;5) maintain biocompatibility; 6) control bioresorption; 7) cast easilyinto complicated anatomical shapes due to greater structural rigidityowing to the nanostructures; and 8) exhibit the mechanical properties ofnative tissues such as articular cartilage.

In one application, the hydrogel/nanostructure composite compositions ofthe invention can be used to repair cartilage tissue. Currentbiologically-based surgical procedures for cartilage repair includeautologous chondrocyte implantation, drilling, abrasion chondroplasty,microfracture, and mosaic arthroplasty. All these procedures treat onlyfocal articular cartilage injuries, and not cartilage denuded jointsurfaces such as seen in severe osteoarthritis and rheumatoid arthritis.Also, they use either cartilage tissue plugs or expanded chondrocytesharvested from the patient to fill cartilage defects. These tissues orchondrocytes are expected to fill the defect by synthesizing entirely denovo material, such as newly synthesized hyaline cartilage, that hasintegrated with existing cartilage matrices and has the biomechanicalproperties of normal cartilage. However, such procedures all promote theformation of a reparative tissue (fibrocartilage) rather than truehyaline cartilage with further mechanical damage to fibrocartilagethought to predispose the joint to osteoarthritis. Furthermore, theavailability of endogenous cartilage as a repair material is quitelimited with its acquisition presenting its own risks and morbidity tothe patient. As evident from the foregoing discussion, the resultinghydrogel/nanostructure compositions disclosed herein present practicalmaterials for promising new therapies in patients suffering fromcartilage degenerative diseases.

As described herein, the present hydrogel/nanostructure compositions canbe prepared having widely varying properties that are suitable for anynumber of synthetic tissue implantation or augmentation, as well asother clinical applications. As already described, the present materialscan be used to repair cartilage defects produced as a result of eitherinjury or disease. Defects due to injury that can be so repaired can besports- or accident-related, and may involve only the superficialcartilage layer, or may include the underlying subchondral bone. Defectsdue to disease which can be repaired using the compositions describedherein include those resulting from osteoarthritis and rheumatoidarthritis. Whether from injury or disease, such defects may be in eithermature or growth plate cartilage. Formulations for hydrogels forsynthetic growth plate cartilage may require the inclusion ofunsubstituted scaffold material to allow for controlled bioresorption ofthe biomaterial during growth.

Another field where the hydrogel/nanostructure compositions describedherein can be useful is the repair, reconstruction or augmentation ofcartilaginous as well as soft tissues of the head and neck. Theavailability of biomaterials for soft tissue augmentation and head andneck reconstruction has remained a fundamental challenge in the field ofplastic and reconstructive surgery. Significant research and investmenthas been undertaken for the development of a material with appropriatebiological compatibility and life span. The outcomes of this researchhave not been promising. When placed in immunocompetent animals thestructural integrity of currently proposed materials has been shown tofail as the framework is absorbed. Furthermore, though conventionalsynthetic materials offer excellent lifespan, they present certainunavoidable pitfalls. For example, silicones have been fraught withconcerns of safety and long-term immune related effects. Syntheticpolymers PTFE (gortex) and silastic offer less tissue reactivity but donot offer tissue integration and can represent long term risks offoreign body infections and extrusion. The materials described in thisapplication will be useful to prepare a synthetic soft-tissue scaffoldmaterial for the augmentation or repair of soft-tissue defects of thehead and neck. In particular, the hydrogel/nanostructure compositions,which are non-inflammatory, non-immunogenic, and which can be preparedhaving the appropriate degree of viscoelasticity (see descriptionherein), could be used as an effective implantable scaffold material.

In addition, the present hydrogel/nanostructure compositions can beused, for example, as a novel, biocompatible and biocompliant materialsto prepare cartilage implants which are frequently used inreconstructive procedures of the head and neck to repair cartilaginousor bony defects secondary to trauma or congenital abnormalities.Applications specific to the ear include otoplasty and auricularreconstruction, which are often undertaken to repair cartilaginousdefects due to trauma, neoplasm (i.e., squamous cell carcinoma, basalcell carcinoma, and melanoma), and congenital defects such as microtia.Applications specific to the nose include cosmetic and reconstructiveprocedures of the nose and nasal septum. Dorsal hump augmentation, tip,shield and spreader grafts are frequently used in cosmetic rhinoplasty.Nasal reconstruction following trauma, neoplasm, autoimmune diseasessuch as Wegeners granulomatosis, or congenital defects require cartilagefor repair. Septal perforations are difficult to manage and often failtreatment. Cartilage grafts would be ideal for these applications, asautologous or donor cartilage is often unavailable. Applicationsspecific to the throat include laryngotracheal reconstruction, which inchildren usually requires harvesting costal cartilage, which is notwithout morbidity. Auricular and septal cartilage is often inadequatefor this application. Synthetic cartilaginous materials prepared fromhydrogels disclosed herein can be synthesized to suit each of theforegoing applications, based on tuning parameters of hydrogel synthesissuch as reagent concentration, substitution and cross-linking rates.Laryngotracheal reconstruction is usually performed for airway narrowingdue to subglottic or tracheal stenosis. The etiology may be traumatic(i.e., intubation trauma, or tracheotomy) or idiopathic. Otherpossibilities include chin and cheek augmentation, and use in ectropionrepair of the lower eyelid, in addition to numerous craniofacialapplications. It should be noted that these applications may not needcartilage with the exacting mechanical properties of articularcartilage. Inclusion of a cell population or bioactive agents may alsobe desirable.

The hydrogel/nanostructure compositions described herein also can beused for repair and narrowing of the nasal cavity, normally followingoverly aggressive surgical resection, to prevent the chronic pooling offluid in the nasal passages that leads to infection and encrustation.Another promising application is in laryngotracheal reconstruction inboth children and adults, as a result of laryngotracheal injury due forexample to intubation during a surgical procedure such as cardiovascularsurgery. Hydrogel/nanostructure compositions as herein described alsocan be used to provide cricoid ring replacements to protect the carotidartery following neck resection for cancer—the composition of theinvention can be placed between the carotid artery and the skin as aprotective barrier for the carotid artery against loss of the skinbarrier. As a protective coating during neuronal repopulation of aresected nerve—often fibrous tissue forms faster than the neuronalrepopulation preventing its eventual formation. Placement of the nerveends within a hydrogel/nanostructure composition of the inventionpre-cast tube could exclude fibrous tissue formation from the site ofrepopulation.

The hydrogel/nanostructure compositions of the invention can also beused for repair of soft tissue defects of any internal or externalorgans. For example, the materials of the invention can be used to forchin and cheek augmentation, and use in ectropion repair of the lowereyelid, in addition to numerous craniofacial applications. For cosmeticand reconstructive purposes in sites other than the head and neck, forexample use as breast implants for breast augmentation, as a woundsealant, for example to fill the void left after removal of lymph nodes(i.e. due to cancer) in the breast or neck, to seal the lymphatics andabate uncontrolled fluid drainage into the resection site that may leadto infection and other complications.

In addition to the above uses, the hydrogel/nanostructure compositionsdescribed herein can be used in other tissue engineering applications toproduce synthetic orthopaedic tissues, including, but not limited to,bone, tendon, ligament, meniscus and intervertebral disc, using similarstrategies and methodologies as described above for the synthesis ofartificial forms of cartilage. The hydrogel/nanostructure compositionsalso can be used to make synthetic non-orthopedic tissues including butnot limited to vocal cord, vitreous, heart valves, liver, pancreas andkidney, using similar strategies and methodologies as described abovefor the synthesis of artificial forms of cartilage.

Another field where the hydrogel/nanostructure compositions disclosedherein can be used is in gastrointestinal applications where it isnecessary to treat or prevent the formation of scar tissue or stricturesin abdominal or gastrointestinal organs. There already are a number ofproducts at various stages of clinical and FDA approval, which generallyare termed “hydrogels,” that are designed or intended to be useful inthe treatment and prevention of scarring and/or stricture formation. Thematerials of the present invention are superior to other known hydrogelsin that the ones disclosed here can include a nanostructure which canprovide support, shape, and strength to hydrogel materials. Thehydrogel/nanostructure compositions disclosed herein can be used insimilar applications as the already known hydrogels are used or intendedto be used, including the following: for treatment of strictures orscarring of the gastrointestinal tract. The treatment involves injectionof the hydrogel material at the site of an anticipated stricture toprevent scarring, or at a site of existing stricture after therapy toenlarge the narrowed GI tract to prevent the stricture from reoccurring.

The materials of the invention can also be used for the treatment ofesophageal strictures. Esophageal strictures are a common complicationof gastroesophageal reflux disease (GERD). GERD is caused by acid, bileand other injurious gastric contents refluxing into the esophagus andinjuring the esophageal lining cells. Approximately 7-23% of GERDpatients develop an esophageal stricture, or fibrous scarring of theesophagus. Esophageal scarring also can be caused by ablative therapiesused to treat Barrett's esophagus. The major complication of suchablative therapies is that the ablative injury extends too deeply intothe esophageal wall and results in an esophageal scar or stricture.Esophageal strictures prevent normal swallowing and are a major cause ofpatient morbidity. The materials described herein may be used to treator prevent esophageal strictures resulting from GERD, Barrett'sesophagus, and esophageal ablative therapies.

The composite materials of the invention may also be used for treatmentof Crohn's disease. Crohn's disease causes strictures or scars thatblock off or narrow the lumen of the bowel, preventing normal bowelfunction. The present materials may be useful to treat or prevent suchstrictures.

The composite materials can also be used in methods for treating primarysclerosing cholangitis (PSC). PSC is a rare disease of the bile ducts ofthe liver. The bile ducts form a branching network within the liver andexit the liver via two main branches that are combined into the commonbile duct which drains the liver and gallbladder of bile into theduodenum. The bile ducts are very narrow in diameter, measuring only upto 2 mm normally at their largest most distal portions, and yet theymust normally drain liters of bile every day from the liver into theduodenum. Any blockage of these ducts can result in a serious conditionknown as jaundice, which allows many toxins and especially hemoglobinbreakdown products to accumulate in the body. PSC is a scarring orstructuring disease of the bile ducts within the liver and in theextrahepatic bile ducts described above that connect the liver to thesmall intestine. The bile duct strictures of PSC may be treated orprevented with the present hydrogel/nanostructure compositions.

The composite materials of the invention can also be used to treatchronic pancreatitis. Chronic pancreatitis is a chronic inflammatorydisease of the pancreas that may be complicated by scars or stricturesof the pancreatic ducts. These strictures block the drainage ofpancreatic juice, which normally must exit the pancreas through a systemof ducts or drainage conduits into the small intestine. The pancreaticjuice contains many digestive enzymes and other elements important tonormal digestion and nutrient absorption. Blockage or narrowing of thepancreatic ducts by chronic pancreatitis can results in severecomplications in which the pancreas autodigests and formslife-threatening abdominal infections and or abscesses. The pancreaticstrictures of chronic pancreatitis may be treated or prevented with thepresent hydrogels.

The presently described compositions may also be used for treatment ofgallstone-induced bile duct and pancreatic duct strictures. Gallstonesare a very common disorder, a principal complication of which is theformation of bile duct and pancreatic duct strictures, which may betreated or prevented with the hydrogels. for treatment of ischemic boweldisease. The intestines are prone to the formation of scars orstrictures when their blood supply is compromised. Compromised bloodflow is called ischemia, and can be caused by many pathologies,including cardiovascular disease, atherosclerosis, hypotension,hypovolemia, renal or hepatic disease-induced hypoalbuminemia,vasculitis, drug-induced disease, and many others. The end stage resultof all of these etiologies can result in intestinal strictures thatblock off the bowel and prevent its normal function. The presenthydrogel/nanostructure composites may be used to treat or preventischemic bowel strictures.

The compositions of the invention may also be used for treatment ofradiation-induced intestinal strictures. Radiation therapy for cancer isassociated with numerous morbidities, important among which isintestinal stricture formation. The present hydrogel composites may beused to treat or prevent radiation-induced intestinal strictures.

In addition to making synthetic tissues or repairing native tissues, thehydrogel/nanostructure composites disclosed here also can be used toprovide a coating for non-biological structures or devices to be used insurgery or otherwise for in vivo implantation, such as surgicalinstruments, or ceramic or metal prostheses. Such a coating wouldprovide a barrier between the non-biologic device material and livingtissue. The role of hydrogels as a barrier for non-biologic devicesincludes, but is not limited to: 1) prevention of absorption ofmacromolecules and/or cells on the surfaces of non-biologic devices,which can lead to protein fouling or thrombosis at the device surface;2) presentation of a non-toxic, non-inflammatory, non-immunogenic,biologically compatible surface for devices made from otherwisenon-biologically compatible materials; 3) compatibility with devicefunction such as diffusion of glucose for a glucose sensor, transmissionof mechanical force for a pressure sensor, or endothelization of avascular graft or stent; 4) enhancement of device function, such asproviding a charge barrier to an existing size barrier in a MEMS basedartificial nephron; 5) incorporation into non-biologic devices of aviable cell population entrapped within an aqueous, physiologicallycompatible environment; and 6) inclusion of drugs or bioactive factorssuch as growth factors, anti-viral agents, antibiotics, or adhesionmolecules designed to encourage vascularization, epithelization orendothelization of the device.

Based on the foregoing, the hydrogel/nanostructure composites of thepresent invention may be used to provide a non-allergenic coating for avariety of implantable devices including an implantable glucose sensorfor management of diabetes. In addition, the hydrogel/nanostructurecomposites may be used to provide: a charge barrier for the developmentof MEMS-based artificial nephrons; an aqueous, physiologicallycompatible environment in which embedded kidney cells such as podocytescan be incorporated into a MEMS-based artificial nephron design; and acoating for implantable MEMS devices designed for a variety of purposesincluding, but not limited to, drug delivery, mechanical sensing, and asa bio-detection system.

The disclosed hydrogel/nanostructure composites, and particularly ahyaluronan-based hydrogel, also may be covalently attached tosilicon-based devices, e.g. through first covalent attachment of theprimary amine of tyramine to the silicon surface to provide ahydroxyphenyl coated surface chemistry. This may use the same chemistryused to bind DNA that has been modified with a free amine to siliconsurfaces. The HA-based hydrogel then is covalently coupled to thehydroxyphenyl coated surface by the same peroxidase driven chemistryused in its preferred cross-linking mode described above.

The hydrogel/nanostructure composites also can be used for coatingnon-biologic cardiovascular devices such as catheters, stents andvascular grafts. These would include devices made from materialsconventionally not used because of their biological incompatibility, butwhich have superior design characteristics to those devices currently inuse. Bioactive factors could be incorporated into the hydrogels topromote endothelization or epithelization of the hydrogel, and thus ofthe implanted device.

Although particular examples and uses for the hydrogel/nanostructurecomposites of the invention have been described herein, such specificuses are not meant to be limiting. The hydrogel/nanostructure compositesof the invention can be used for any application generally used forknown hydrogels, and in particular, are useful for the repair and/orregeneration of soft tissue anywhere in the body.

Reference will now be made to the drawings wherein like referencenumerals identify similar structural features or aspects of the subjectdisclosure. For purposes of explanation and illustration, and notlimitation, an illustrative view of an embodiment of a biodegradablecomposite in accordance with the disclosure is shown in FIG. 1A and isdesignated generally by reference character 100. The systems and methodsdescribed herein can be used to enhance healing of soft tissue defects.

Referring generally to FIGS. 1A-1D, the biodegradable composite 100 caninclude a nanofiber 101 reinforced gel 103 that combines the advantagesof both gel 103 and nanofibers 101. The gel 103 can include any suitablematerial, such as, but not limited to, hydrogel. The nanofibers 101 canbe made of any suitable nanomaterial, e.g., polycaprolactone (PCL) orany other suitable material, and can take any suitable shape and/orsize. The composite 100 includes high porosity (e.g., to mediate celladhesion and migration) while maintaining sufficient mechanicalproperties (e.g., to maintain integrity and tissue support).

In at least some embodiments, the nanofibers 101 are covalentlyconjugated to the hydrogel 103 forming one or more polymer chains.Covalent attachment of hydrogels 103 to the nanofibers 101 can result ina material with a combined set of ideal properties superior to theconstituent materials used alone or as a simple blend.

FIG. 2A depicts stress-strain curves of an embodiment of the compositeof FIG. 1 plotted against HA Hydrogel alone, revealing improved elasticmodulus compared to hydrogel at the same crosslinking density. As shown,the elastic modulus of the tested composite 100 (4.5 mg/ml HA, 10 mg/mlPEG-DA, 6.75 mg/ml PCL fibers) was 750 Pa, and hydrogel alone at thesame density was 320 Pa. FIG. 2B depicts a fatigue test showing that thecomposite as illustrated in FIG. 1 retains similar degree of robustnessof mechanical integrity compared to regular hydrogel

Referring to FIG. 3A-3B, the composite 100 was shown to supportadipose-tissue derived stem cell (ASC) migration. GFP-labelled ASCs fromliposuction aspirates were grown into spheroids and then seeded intocomposite or hydrogel.

FIGS. 3A and 3B show fluorescence and overlay (FIG. 3A) with phasecontrast images (FIG. 3B) of ASCs cultured in nanofiber-HA hydrogelcomposite for 4 days. The cells migrated outwards with extended longprocesses and trajectories. In contrast, ASCs cultured in HA hydrogelalone shown in FIGS. 3C and 3D did not show significant cell migration.

FIGS. 4A and 4B show a fluorescence image and overlay (FIG. 4A) withphase contrast image (FIG. 4B) contrasting ASCs migrating from spheroidsalong aligned 650-nm nanofibers 101, showing their strong migratoryresponse to the presence of nanofibers 101.

Example 1: Preparation of a Composite Surgical Scaffold Device

Nanofibers were produced by electrospinning PCL (polycaprolactone, 80 kfrom Sigma Aldrich). The nanofibers were spun into a random mesh. Thespinning parameters were a 10% wt solution of PCL in 90%/1% w/w DCM-DMF,at a flow rate of 0.6 ml/h through a 27 gauge blunt needle 15 cm fromthe target metal plate. The needle voltage was +10 kV, with the targetplate was negatively biased with a voltage of −3 kV. One mL of solutionwas spun per round for each of the nanofiber sheet.

The fibers were then functionalized with a multistep process. Briefly,the fibers were plasma-treated to have reactive groups on the fibersurface, to which acrylic acid was conjugated by UV photoinitiation. Theacrylate groups were then reacted with EDC and diazimine to form primaryamines. These amines could then be reacted with SMCC to attach maleimidegroups, which could readily react with the thiol groups in the hydrogel.

A composite surgical mesh was prepared using a composite gel formulationwith 5.4 mg/mL of thiolated hyaluronic acid (220 KDa, thiolation degreeof 25%) and 5.4 mg/mL of PEG-diacrylate (PEG-DA) with 10 mg/mL offunctionalized dispersed nanofibers. The polypropylene surgical meshused was Ethicon Prolene Soft (Product code SPMH). The meshes werecleaned with serial ethanol soaks and allowed to dry in the biosafetycabinet before use. The 1×2 cm rectangles of mesh were placed into thebottom of the 2.5×4.5 cm Teflon molds, two per mold.

An aliquot of 500 μL of composite was pipetted on into each mold (forboth meshes), then a piece of plastic was placed over the meshes, andpressed down to spread out the composite. The meshes were allowed to gelovernight in 37° C. incubator. The gelled meshes were removed andlyophilized as a final product (see FIG. 21). The mesh could then berehydrated prior to use.

In certain embodiments, the functionalized-fiber mesh was cut intosections of 60 mg or less. A 60 mg sample is soaked in ethanol, and thenadded to the ceramic mortar that has been partially filled with liquidnitrogen. The fiber sample will become very rigid. Keeping the samplecool enough to maintain rigidity, the fiber sheet is cut into ˜5 mm×5 mmsections with scissors. When the full sheet has been cut, the fibers areground with the mortar and pestle for ˜20 min, keeping the mortarpartially full with liquid nitrogen. The fiber slurry is then pouredinto ethanol. About 1 mg of surfactant is added to the slurry to helpprevent fiber entanglement. The suspension is centrifuged for min at 300G, and the supernatant is discarded. The fibers are allowed to dryovernight. The fibers are then weighed into a secondary centrifuge tube,so that a precise concentration of fibers can be suspended. The fibersare then soaked in ethanol to sterilize, centrifuged, had thesupernatant discarded, and allowed to dry overnight in a biosafetycabinet. The fibers are then resuspended to the desired concentration indeionized water, usually 15 mg/mL.

To form the hydrogel composite, 1 mL of the fiber-suspension is used torehydrate 1 vial of HA-SH, resulting in a solution of 15 mg/mL fibersand 10 mg/mL of hyaluronic acid. To 900 μL of this solution, 100 μL of10% PEG-DA stock solution is added, to give a final concentration of13.6 mg/mL fiber, 9 mg/mL HA-SH, and 10 mg/mL PEG-DA. This is theformulation for the initial in vivo examples, but other formulationshave been made by varying the constituent concentrations.

The resulting composites were milky white in color (FIG. 1C), as opposedto the transparent hydrogels without the fibers. The composite gelsmaintained their shape and had good handleability, while thehydrogel-alone group was more prone to tearing. The fibers in thehydrogel were disperse and ranged in length from tens to hundreds ofmicrons (FIG. 1B). A SEM image of cross-section of a fractured,lyophilized sample composite shows the close association between thefibers and hydrogel component, as well as the high density of dispersedfibers (FIG. 1D).

Materials and Methods

Thiolated hyaluronic acid (HA) was purchased from ESI BIO (Alameda,Calif.). Poly(ethylene glycol) diacrylates was purchased from LaysanBio, Inc (Arab, Ala.). The followings were obtained from Sigma;poly(ε-caprolactone), ethylamino-maleimide, acrylic acid, Toluidine blueO, N-hydroxysuccinimide (NHS), cysteine, bovine serum albumin (BSA),acetic acid and Triton™ X-100. Dulbecco's modified eagle medium (DMEM),fetal bovine serum (FBS), penicillin/streptomycin, Alexa Fluor® 568Phalloidin and 4′,6-diamidino-2-phenylindole (DAPI) were purchased fromInvitrogen Life Technologies. Ethyl(dimethylaminopropyl) carbodiimide(EDC) was obtained from AnaSpec, Inc. (Fremont, Calif.). All otherchemicals and reagents were of analytical grades.

Electrospinning of PCL Nanofibers for Rheology Experiments:

To fabricate two different diameters of PCL fibers, 11.0 and 8.5% (w/v)PCL solution were prepared in a mixture of dichloromethane anddimethylformamide (9:1, v/v) and a mixture of chloroform and methanol(3:1, v/v), respectively. Each homogenous PCL solution was loaded asyringe with a metallic needle of 27 G. Then, electrospinning wasperformed with following parameters; 1.0 ml/h of a feeding rate, 15 kVof an applied positive voltage for a metallic needle, and 12 cm of adistance between the end of a needle to a ground. Morphology of fiberswas observed using a field-emission scanning electron microscope (FESEM,JEOL 6700F) and a diameter of fibers was measured with FESEM imagesusing ImageJ software (US National Institutes of Health, Bethesda, Md.).

Electrospinning for In Vivo Composites:

Spinning conditions: 16% w/v PCL (95% 45.000 Mn PCL, 5% 80,000 Mn PCL,both from Sigma) in a solvent mixture of dichloromethane anddimethylformamide (9:1, w/w). The fibers were spun at a rate of 5.25ml/h through a blunt 27 gauge needle separated 10 cm from the face ofthe grounded wheel, spinning at 1000 rpm. The applied voltage was 15 kVand the electrospinning pump was rastered back and forth across the 85mm travel distance for 140 passes at 2 mm/sec (about 4 h). The fibersheet was then cut into 14 cm-diameter individual sheets forfunctionalization.

Preparation of Surface-Functionalized Fibers with MAL:

To surface-functionalize on fibers with MAL, a surface of fibers wasinduced carboxyl groups by grafting poly(acrylic acid) (PAA) accordingto the literature with a minor modification [Interface Focus 2011, 1,725-733]. Briefly, fibers were plasma-treated under 280 mmHg with oxygenatmosphere at room temperature for 10 min to induce free radicals on asurface of fibers. Then 70 mg of fibers in 10 ml of 3 or 10% (v/v)acrylic acid solution in 0.5 mM NaIO3 was exposed to UV (36 mW/cm²,DYMAX Light Curing Systems 5000 Flood, Torrington, Conn.) for 90 s forphoto-polymerization of PAA on fibers surface (PAA-fibers). Afterincubating PAA-fibers at room temperature for 20 min, PAA-fibers werewashed with 20 ml of deionized water three times to remove unreactedacrylic acid. After completely air-drying PAA-fibers, a density ofcarboxyl groups on PAA-fibers were determined by toluidine blue O (TBO)assay with the assumption that TBO interacts with a carboxyl group onfibers at 1:1 of molar ratio [J Biomed Mater Res 2003, 67, 1093-1104].Briefly, PAA-fibers (1×1 cm²) were completely immersed in 1 ml of 0.5 mMTBO solution in 0.1 mM of NaOH (pH 10) after soaking 20 μl of 50% (v/v)ethanol and reacted with gentle shaking at room temperature for 5 h.After washing them with 0.1 mM NaOH (pH 10), adsorbed TBO on a surfaceof PAA-fibers was desorbed using 1 ml of 50% (v/v) acetic acid withvigorous shaking at room temperature for 1 h. Then an optical density ofsupernatant was measured at 633 nm using a microplate reader (BioTeckSynergy2, Winooski, Vt.). TBO in 50% (v/v) acetic acid was used as astandard.

PAA-fibers were ground to prepare fiber fragments using a cryogenic mill(Freezer/Mill 6770, SPEX SamplePrep, Metuchen, N.J.) with followingparameters; 10 cycles of 1 min for milling and 3 min for cooling inliquid nitrogen. After collecting PAA-fiber fragments into a 50-mlconical tube, PAA-fiber fragments were completely dispersed in 10 ml ofa mixture of isopropylacohol and distilled water (1:1, v/v) to modifywith aminoethyl-MAL on a surface of fibers. Briefly, PAA-fibers wereadded NHS and EDC to activate carboxyl groups of PAA on fibers. A molarratio of carboxyl group to NHS and EDC was 1 to 4 and 4, respectively.The activation was performed with gently shaking at room temperature.After 1 h, aminoethyl-MAL was added into the carboxyl groups-activatedfibers with 1 to 2 of molar ratio of carboxyl groups to aminoethyl-MAL.Then the reaction was performed with gently shaking at room temperaturefor 12 h. Surface-functionalized fibers with MAL were lyophilized afterwashing with distilled water three times. Here, a density of MAL onfibers was on the assumption that all of carboxyl groups on a surface offibers were completely substituted by MAL.

Preparation of Fiber-HA Hydrogel Composites:

For preparing a fiber-HA hydrogel composite, thiolated HA and PEGDA werecompletely dissolved in PBS (pH 7.4) to the desired concentration of12.5 mg/mL and 100 mg/mL, respectively. MAL-fibers with the desiredconcentration of 25 mg/mL were completely dispersed in PBS (pH 7.4). Thesuspension of nanofibers, HA, PEG-DA, and PBS are then serially added toreach the formulation's desired final concentration. After homogenousmixing the composite precursor solution, for rheological studies, 100 μLof the composite precursor solution was poured into a mold (diameter=8mm) and incubated at 37° C. for 2 h for gelation. For compressionstudies, 200 μL of precursor solution is added to a cylindrical Teflonmold (diameter=6.35 mm, h=6.35 mm) and incubated as above. To observemorphology of cross-section of a fiber-HA hydrogel composite and HAhydrogel using FESEM, a composite and HA hydrogel were dehydrated byserial ethanol washing (10 min each at 50%, 70%, 80%, 90%, 100%, and100% Ethanol) before either critical point drying (Samdri-795, Tousimis,Rockvillle, Md.) or chemical drying (HDMS). The samples werefreeze-fractured in liquid nitrogen to reveal the internal porestructure. The structure was sputter coated with a 10-nm layer ofplatinum (Hummer 6.2 Sputter System, Anatech UDA, Hayward, Calif.), thenimaged with a field-emission SEM (JEOL 6700F, Tokyo Japan).

For preparation of the composites for the in vivo animal studies. thethiolated HA was reconstituted to 12.5 mg/mL in PBS. The PEG-DA wasdissolved to 100 mg/mL in PBS. The MAL-fibers were resuspended to 25mg/mL in sterile PBS. The fibers were first combined with the HAsolution and allowed to react for 10 min before being combined with thePEG-DA to obtain the desired final concentrations. The suspension wasthen immediately pipetted into the cylindrical Teflon molds(McMaster-Carr, Robbinsville, N.J.), with 300 μL into cylindrical molds11.125 mm in diameter and 3 mm in height for the in vivo samples. Thegels were then placed into the 37° C. incubator to gel overnight.

To confirm the effect of interfacial bonding between thiol groups of HAand MAL on fibers, MAL on fibers was quenched using cysteine forpreparing a quenched fiber-HA hydrogel composite. Briefly, 1 mg offibers was dispersed in 1 ml of cysteine solution in PBS (pH 8.0) then amolar ratio of MAL to cysteine was 1 to 2. After quenching the MAL withgentle shaking at room temperature for 12 h, MAL-quenched fibers werewashed with 1 ml of distilled water five times to remove unreactedcysteine and lyophilized.

Mechanical Properties of Fibers-HA Hydrogel Composites:

Compressive test. The hydrogel precursor suspension was pipetted intothe cylindrical Teflon molds (McMaster-Carr, Robbinsville, N.J.), with200 μL into cylindrical molds 6.35 mm in diameter and 6.35 mm in heightfor compression testing. The gels were then placed into the 37° C.incubator to gel overnight. The gels were removed from their molds andimmediately tested via unconfined uniaxial compression between twoparallel plates with the Endura TEC mechanical tester ELF 3200 Series,BOSE ElectroForce, Eden Prairie, Minn.). The samples were compressed to50% strain, with the elastic modulus determined from the slope of thelinear portion of the stress-strain curve from 10% to 20% strain. Thesamples were tested three times each, and three samples were tested pergroup for determining the average compressive modulus. To measurecompressive modulus of rehydrated fiber-HA hydrogel composites, thecomposites were lyophilized and rehydrated with 1 ml of PBS (pH 7.4) at37° C. for 24 h. For fatigue-testing, the compression samples wererepeatedly cycled from 0% to 25% strain at 0.1 Hz.

Rheological test. Shear storage modulus (G′) of various fiber-HAcomposites were measured using an oscillating rheometer (ARES-G2Rheometer, TA Instruments, New Castle, Del.) with a parallel plate (ϕ=8mm). Oscillatory frequency sweep was employed to monitor variation of G′from 1 Hz to 10 Hz with constant strain of 10%.

Migration of hASCs in Fiber-HA Hydrogel Composites:

Human adipose-derived stem cells (hASCs) were cultured in high glucoseDMEM containing 10% of FBS, 1% of penicillin/streptomycin, and 1 ng/mlof bFGF. The culture medium was exchanged three times per a week foroptimal growth. To prepare hASC spheroids, 50 μl of hASCs solution(5.6×10⁵ cells/ml) was poured into a casted micro-molded agarose gel(MicroTissues® 3D Petri Dish® micro-mold spheroids, 96-holes) to preparehASCs spheroids and incubated with gently shaking at 37° C. for 24 h.

HA and PEGDA were completely dissolved in PBS (pH 7.4) with finalconcentration of 4.5 and 2.5 mg/ml for HA and 5.0 mg/ml for PEGDA.Fibers pre-wetted with 20 μl of 50% (v/v) ethanol were completelydispersed in PEGDA with final concentration of 10.0 mg/ml, then HA addedinto a mixture of fibers and PEGDA. 30 μl of composite precursorsolution was poured into each well of a 96-well tissue culture plate andincubated to crosslink at 37° C. for 1 h for avoiding to reach hASCsspheroids on a surface of tissue culture plate. Then, 50 μl of compositeprecursor solution with 3˜5 of hASCs spheroids was poured into the eachwell. After crosslinking at 37° C. for 1 h, 200 μl of fresh media wereadded into the each well and the media were exchanged every a couple ofdays. To observe migrated cells from hASCs spheroids inside thecomposites, F-actin and nuclei of hASCs were stained with Alexa Flour®568 Phalloidin and DAPI, respectively. Briefly, after 4 days ofcultivation, the composites with hASCs spheroids were fixed with 100 μlof 4% (v/v) paraformamide at room temperature for overnight. Then, afterwashing three times with PBS (pH 7.4), the composites were incubatedwith 100 μl of 1% (w/v) BSA in PBS to inhibit non-specific staining at4° C. for overnight and washed three times with PBS. Subsequently, thecomposites were incubated with 100 μl of 0.1% (v/v) Triton-X 100 in PBSat room temperature for 1 h. After washing three times with PBS, 100 μlof 160 nM Alexa Fluor® 568 Phalloidin was added into each composites andincubated at room temperature for 4 h. Then, after removing thesupernatant, the composites were incubated with 100 μl of 0.5 μg/ml DAPIat room temperature for 1 h. After washing three times with PBS, themigrated hASCs were observed using confocal laser scanning microscope(CLSM, Carl Zeiss LSM780, Germany) at ex. 561 nm and em. 570-600 nm forAlexa Fluor® 568 Phalloidin, and ex. 405 nm and em. 385-420 nm for DAPI.

Performance of a Fiber-Hydrogel Composite In Vivo:

The thiolated HA was reconstituted to 12.5 mg/mL in PBS. The PEG-DA wasdissolved to 100 mg/mL in PBS. The MAL-fibers were resuspended to 25mg/mL in sterile PBS. The fibers were first combined with the HAsolution and allowed to react for 10 min before being combined with thePEG-DA to obtain the desired final concentrations. The suspension wasthen immediately pipetted into the cylindrical Teflon molds(McMaster-Carr, Robbinsville, N.J.), with 300 μL into cylindrical molds11.125 mm in diameter and 3 mm in height. The gels were then placed intothe 37° C. incubator to gel overnight. The two formulations wereselected so as to match the 2 kPa stiffness of fat tissue. The HA-aloneformulation was 10 mg/mL PEG-DA and 9 mg/mL HA-SH, and the HA-fibercomposite formulation was 5 mg/mL PEG-DA, 5 mg/mL HA-SH, and 12.5 mg/mLdispersed nanofibers.

To study the biocompatibility of the composite nanomaterial scaffolds,they were implanted under the inguinal fat pads of Sprague-Dawley ratsand observed for varying lengths of time. Under volatile anesthesia, a 1cm incision was made just proximal to the inguinal crease bilaterally.Following blunt dissection of subcutaneous tissues, the inguinal fat padwas exposed. It was elevated with meticulous hemostasis usingelectrocautery and with careful preservation of feeding vessels.Scaffolds were implanted under the fat pad on the right side of theanimal. The left side received no implant and served as sham surgerycontrol. Both sides were closed in a standard layered fashion. Animalswere observed for 7, 14, 30, and 90 days. At timepoints for collection,animals were sacrificed and the inguinal fat pad with and withoutscaffolds was exposed and fixed in 4% PFA. The specimens were imbeddedand sectioned for standard hematoxylin and eosin staining.

Statistical Analysis

All the results are expressed in mean values and the standard deviation.The statistical significance between a pair of groups was determined byconducting a One Way ANOVA with SigmaPlot 12.0 software (SPSS); a valueof p<0.05 was considered statistically significant.

Any other suitable method for making embodiments of the composite 100 asdisclosed herein are contemplated herein.

Example 2: Compression Test of the Nanofiber-Hydrogel Composite

For compression testing, the fiber-hydrogel samples were formed ascylinders 8.5 mm in diameter and ˜4 mm in height, allowed to setovernight in molds at 37° C. The elastic moduli were determined viacompression testing with a Bose EnduraTEC ELF 3200 (Eden Prairie,Minn.). The sample underwent uniaxial compression between two parallelplates, compressed to 50% Strain. The elastic moduli were determined bymeasuring the slope of the initial linear region. Two sample groups weretested, with the same hydrogel formulations, with and without fibers.The hydrogel-only sample was formed with 4.5 mg/mL of thiolatedhyaluronic acid (Gylcosan Glycosil) and 10 mg/mL PEG-DA(polyethyl-glycol diacrylate, molecular weight 3350). The fiber-hydrogelcomposite group had the same hydrogel concentrations, but additionallyhas 6.75 mg/mL PCL nanofibers that have a surface functionalized withmaleimide groups that can readily react with the thiolated hyaluronicacid.

Representative stress-strain traces can be seen in FIG. 2A. Thehydrogel-only group had an elastic modulus of 320 Pa, while thefiber-hydrogel composite had a higher modulus of 750 Pa. Thefiber-hydrogel composite's increased stiffness can be seen in the higherstress values at every strain value. The presence of functionalizednanofibers greatly increased the strength and stiffness of the material.Thus, the overall structure of the composite can have a stiffnessmatched to the target tissue, while the hydrogel component can have alower crosslinking density than the density that would be needed toachieve the same stiffness without the benefit of nanofibers. Thisshould result in a better cellular response for a given implantstiffness.

The sample groups were then tested via repeated compression to 25%Strain (20 cycles) at 0.1 Hz. Representative traces can be seen in FIG.2B. This shows that the hydrogels and composites can tolerate repeatedcompression, and that the composite is persistently stiffer than thefiberless group.

Example 3: Cell-Materials Interaction

To test for the cellular response to the composite hydrogel, themigratory potential of adipose-derived stem cells (ASCs) was tested invarying formulations of hydrogels, with and without fibers.

ASCs were transfected to express GFP, then formed into spheroid clustersby seeding the cells overnight in alginate molds made by Microtis suesmolds. The cells were seeded as spheroids to better evaluate cellmotility, as the spheroids are a distinct point source from whichmigrating cells can be easily measured. The spheroids were mixed in tothe hydrogel before being pipetted into a 96-well plate and beingallowed to set. The cells were then imaged over the next several days toobserve their migration. The cells were able to migrate progressivelyfurther as the concentrations of hyaluronic acid and PEG-DA werelowered, due to the respectively increasing pore sizes. At the samehydrogel densities (4.5 mg/mL hyaluronic acid and 2.5 mg/mL PEG-DA),cells were better able to migrate in samples with disperse nanofibers(12 mg/mL, FIGS. 3A and 3B) than without (as shown in FIGS. 3C and 3D).This indicates that the presence of functionalized nanofibers not onlyimproved the mechanical properties of the nanofibers, but also can aidin improving cell migration.

To clearly demonstrate that the ASCs were strongly influenced by thepresence of nanofibers, ASC spheroids were cultured on aligned nanofibersheets, without hydrogel. After 96 hours, the cells (green in FIGS. 3Cand 3D) clearly migrated out of the spheroid along the same axis of thealigned nanofibers (shown in FIG. 3D.)

Example 4: Tissue Compatibility of the Nanofiber-Hydrogel Composite

To study the biocompatibility of the composite nanomaterial scaffolds,they were implanted under the inguinal fat pads of Sprague-Dawley ratsand observed for varying lengths of time. Under volatile anesthesia, a 1cm incision was made just proximal to the inguinal crease bilaterally.

FIG. 5A is a photograph showing appearance of nanofiber-hydrogelcomposite in situ under rat inguinal fat pad. FIG. 5B shows H&E stainingimages of sections from tissues around the composite harvested at 2weeks after implantation. Fronds of eosinophilic, dark pink stainedmesenchymal cells are shown migrating into the nanomaterial (stained inlight pink).

FIG. 5C shows H& E staining images of tissue sections collected fromcomposite-tissue interface at 4 weeks, showing cell infiltration.Mesenchymal tissue surrounding the site of implantation stains dark pinkwith eosin. The nanomaterial appears light pink. Infiltrating pinkmesenchymal cells can be seen at the interface as well as putativeadipocytes with clear round vacuoles.

Following blunt dissection of subcutaneous tissues, the inguinal fat padwas exposed. It was elevated with meticulous hemostasis usingelectrocautery and with careful preservation of feeding vessels.Scaffolds were implanted under the fat pad on the right side of theanimal. The left side received no implant and served as sham surgerycontrol. Both sides were closed in a standard layered fashion. Animalswere observed for 2, 4, and 6 weeks. At timepoints for collection,animals were sacrificed and the inguinal fat pad with and withoutscaffolds was exposed and fixed in 4% PFA. The specimens were imbeddedand sectioned for standard hematoxylin and eosin staining. At earlytimepoints (2 weeks), mesenchymal cells from the wound bed were foundinfiltrating the material suggesting that the material has sufficientporosity to enable native cellular ingrowth (dark pink staining in FIG.5B).

Importantly cellular in-growth was achieved even in the absence ofexogenous growth factors. The presence of cells infiltrating thematerial rather than merely surrounding it, distinguishes this compositenanomaterial from other alloplastic materials in current use. The lattermaterials are walled off by fibrous capsule and are therefore lessdesirable for soft tissue reconstruction. At later timepoints (4 weeks),cellular ingrowth is even more apparent with the appearance of vacuolarareas that may represent nascent adipocyte differentiation (dark pinkstaining and clear circles in FIG. 5C).

Example 5: Design of a Fiber-HA Hydrogel Composite

The fibers could form the fibrous architecture that can often be seen inthe native extracellular matrix, aiding cell migration and reinforcingthe initially-low mechanical properties of the hydrogel. By introducinginterfacial bonding between the hydrogel and fibers (FIG. 6A,FIG. 6B),the composite is strengthened without decreasing the average pore sizeand porosity (FIG. 6) that would significantly hinder cell migration. Itwas also expected that the mechanical properties could be tuned bycontrolling the density of the interfacial bonding between the hydrogeland the surface of fibers. Here, surface-functionalized fibers wereprepared with maleimide (MAL) to introduce the interfacial bonding withthiolated hyaluronic acid (HA-SH) (FIG. 6). The surface of electrospunpoly(ε-caprolactone) (PCL) fibers was treated with O₂ plasma to inducefree-radicals onto its surface before grafting poly(acrylic acid) (PAA).The carboxyl groups was activated by coupling reagents, NHS and EDC,then N-(2-aminoethyl)maleimide was reacted to the activated carboxylgroups (FIG. 13). Subsequently, MAL-functionalized fibers wereintroduced to hydrogel precursor solution composed of HA-SH and PEGDAfor fabricating a fiber-hydrogel composite. The thiol groups of the HAwere employed to form a gel by reacting with both the MAL groups on thefibers and the DA groups of the PEG linker. Interestingly, across-section of a fiber-hydrogel composite showed a fibrous 3Dstructure with a high porosity (FIG. 6), compared to a cross-section ofHA hydrogel with a similar crosslinking density. The resultingcomposites showed even distribution of nanofibers across both the widthand height of the composite, enabling isotropic reinforcement. Also, arehydrated fiber-HA hydrogel composite showed 99.34% of volume recoveryafter lyophilization while HA hydrogel showed 70.17% of volume recovery(FIG. 6D).

Example 6: Compressive Modulus of a Fiber-HA Hydrogel Composite

First, the composite was verified to possess its maximal stiffness(under shear) when the reactive groups were equal on a molar basis. Thethiol groups on the HA can react with either the MAL groups on thenanofibers or the acrylate groups on the PEG-DA, so when the molar ratioof SH to (DA+MAL) was approximately 1 to 1, the gels showed an optimalshear storage modulus. Therefore, this ratio was maintained for all ofthe subsequent studies. The gels underwent unconfined compressiontesting to evaluate the elastic modulus of HA hydrogel and fiber-HAhydrogel composites (FIG. 7). The reinforcing effect of thefunctionalized nanofibers can be seen in the compressive stress whenstrained to 50% (FIG. 7A). The compressive stress was 3.1-fold greaterin the 1.0-μm fiber group than the hydrogel-only group, showing theeffect of mechanical reinforcement. The 286-nm fiber group showed evenmore pronounced reinforcement effect with a compressive stress of4.2-fold higher at the 50% strain. Interestingly, the stiffening effectof the 286-nm fibers was greatly reduced to only 1.3-fold over thehydrogel when the maleimide groups were quenched prior to gelation,confirming that the interfacial bonding of the fiber to the hydrogel iscrucial to the reinforcement effect of the functionalized fibers.Moreover, when the 286-nm fibers were not functionalized before formingthe composite, the reinforcement effect was disappeared, resulting incomposites barely stiffer than the hydrogel alone. The samereinforcement effect can be seen when formulating stiffer gels byformulating composites with higher concentrations of HA and PEG-DA (FIG.7). The interfacial bonding also shows a dose-response in its stiffeningof the composite gel, as adding progressively more maleimide groups tothe nanofiber surface results in progressively stiffer materialsproviding more evidence of the importance of the interfacial bonding.The composites were also tested for changes in mechanical propertiesbefore and after dehydration and rehydration. The gels, with and withoutfunctionalized nanofibers of two different maleimide densities, weremechanically tested under compression. The gels were then lyophilized,then allowed to rehydrate fully and tested for compression again. Allsamples maintained their stiffnesses after rehydration, indicating thatthe composites may be suitable for use clinically as a lyophilizedproduct. While the HA-alone gel seemingly maintained its stiffness, thegel itself had compacted significantly during thedehydration-rehydration process, unlike the fiber-containing groups. Thecomposite gels were also subjected to cyclic loading to test forfatigue-effects, with representative traces shown in FIG. 10. Withrepeated loading to 25% strain, the composite gels maintained theirstiffnesses over time and were consistently stiffer than the hydrogelalone.

Example 7: Shear Storage Modulus of a Fiber-HA Hydrogel Composite

In addition to the higher compression modulus, the Fiber-HA hydrogelcomposites showed a significantly higher shear storage modulus than theHA hydrogel alone (FIG. 8A). The shear storage modulus of a compositewith 286-nm fibers was higher than that of a composite with 686-nmfibers (FIG. 8C). It was also confirmed that the shear storage modulusof the composites increased by increasing the maleimide surface densityon the 286-nm fibers, similar to the modulus under compression testing(FIG. 8D). By introducing fibers with 62 nmol/mg MAL on its surface, thecomposite showed a 1.3-fold increase in its shear storage moduluscompared to that of the HA hydrogel alone. Moreover, the shear storagemodulus of a composite with 147 nmol/mg MAL on its fibers was increased1.8-fold over the modulus of the 62 nmol/mg MAL group, showing a cleardose response to the corresponding 2.4-fold increase in the MAL surfacedensity on the fibers. When the MAL groups on the fibers were quenchedprior to gelation, the shear storage modulus correspondingly decreasedcompared to that of the unquenched fibers, similarly to what was seen inthe compression testing. Additionally, the shear storage modulus of thecomposites was maintained when the frequency increased to 10 Hz whileboth the HA hydrogel alone and the composite with quenched fibers showeddiminishing shear storage moduli at 10 Hz than those at 1 Hz. The shearstorage modulus of the composites was increased with increasing MALsurface density on fibers regardless of surface area (diameter) offibers, indicating that the previously observed effect of fiber diameteron stiffness may have been a function of maleimide density (FIG. 8D). Alinear regression was obtained from the correlation between the MALsurface density and shear storage modulus with R2=0.93. Moreover, thecomposites showed a dose response to fiber loading, as the shear storagemodulus of the composites increased with an increasing weight ratio offunctionalized fibers to hydrogel components (FIG. 9).

Example 8: Cell Migration in a Fiber-HA Hydrogel Composite In Vitro

It was hypothesized that the fiber-HA hydrogel composite enhanced cellmigration compared to HA hydrogel because of (i) a higher porosity ofthe composite with a larger pore size, providing a spatiality for cellmigration when they have the same mechanical properties and (ii) anECM-mimicked fibrous architecture in the composite, allowing tointrinsically guide cell migration. Therefore, for demonstrating thecurrent hypothesis, spheroids of human adipose-derived stem cells(hASCs) as a model cell were seeded and a mimicked tissue chunk insideHA hydrogel and composites, then the hASCs spheroids were cultured for27 days (FIG. 11). ASCs were chosen due to their presence in fat tissuesand importance in both angiogenesis and adipocyte formation. Althoughthe composites have the similar Young's modulus, 1.9 kPa, to the HAhydrogel, the pore size of the composites is 2.08-fold bigger than thatof HA hydrogel (FIG. 16). Hence, it was clearly observed that hASCsmigrated 3-dimensionally inside the composites (FIG. 11B-11E) becausethe bigger pores could accommodate to migrate the cells, while hASCsmaintained their spheroid shape without any cell migration in HAhydrogel (FIG. 11A). In particular, the cell migration was magnificentlyenhanced when the fibers were modified with the cell adhesion peptide,RGD, for the composite (FIG. 11C). However, in the in vivo setting,diffusion of factors into the composite from the local milieu shouldprovide additional adhesive cues, lessening this difference. In someinstances, partial fibers slightly formed a cluster during gelation dueto hydrophobic interaction between PCL fibers, and it was observedbodies of cells preferentially grabbing the fibers clusters inside thecomposites (FIGS. 11D and 11E). Furthermore, at the same HA and PEG-DAconcentrations (FIG. 19), the composites showed enhanced cell migrationas compared to the fiberless group, showing that the nanofibersthemselves could intrinsically help guide cell migration regardless ofthe porosity.

Example 9: Tissue Response and Host Tissue Infiltration

To determine the therapeutic potential of these composite implants, thecomposite implants were tested in vivo in a rat fatpad model. Theformulations of the implant groups were formulated to achieve the sameinitial 2 kPa stiffness as the composite gel and the target adiposetissue. Thusly, the formulation of the HA-gel alone implant had a higherconcentration of both thiolated HA and PEG-DA to match the stiffness ofthe fiber-composite group. Despite the higher concentrations, theHA-alone implants were unable to maintain their shape and volume overthe course of the study. Under gross observation after 4 weeks, theHA-alone implants were stretched out and significantly smaller involume. Considering their gross appearance and their histological lackof infiltration, the HA-alone system cannot be optimized to be able toencourage cell infiltration and maintain a predetermined shape. Thefiber-gel composite implants, however, well maintained their originalshape under gross observation after 90 days in vivo. Remarkably,however, under histological observation the composites had been sothoroughly infiltrated that the border between implant and native tissuehad become difficult to determine.

A soft tissue defect model in Lewis rats has been developed, where theinguinal fat pad is exposed and elevated using microsurgical techniquesand the pre-shaped composites are placed underneath. This well-definedmodel is ideal to address all elements of Aim 3 hypothesis and the scaleamenable to the R21 study. Even though this does not directlydemonstrate the ability of such a composite for restoring large defects,it will establish the proof-of-principle and confirm all essentialfunctionalities of the composite design, and lay the foundation forlarger animal model to test the large defect restoration in moreclinically relevant models.

In a pilot study, PCL nanofiber-HA hydrogel composites and HA hydrogelswere implanted with similar moduli under the inguinal fat pad of 8-12week old male Lewis rats (n=3 per time point). Both the HA hydrogel andcomposite groups showed good tissue compatibility at days 14 and 30after transplantation (FIG. 12, POD 14, similar observations at POD 30.POD=Post-Operative Date). Histology at POD 30 did not show higher levelof inflammatory response than sham surgery group. H&E and Masson'strichrome staining showed septation and cellular infiltration by nativefat through the composite, capillary formation around the perimeter, andregeneration of glandular as well as adipocyte portions of native fat(FIG. 12). HA hydrogel control on the other hand, lacked cellularinfiltration and formed a thin sheet of fibrotic tissue and foreign bodyresponse. This HA hydrogel was prepared with 2 kPa to ensure sufficientmechanical property. This result highlights the importance of porosityof the scaffold for cell infiltration.

At an early time point (2 weeks), mesenchymal cells from the wound bedwere found infiltrating the material suggesting that the material hassufficient porosity to enable native cellular ingrowth (dark pinkstaining in FIG. 12). Importantly, cellular ingrowth was achieved evenin the absence of exogenous growth factors. The presence of cellsinfiltrating the material rather than merely surrounding it,distinguishes this composite nanomaterial from other alloplasticmaterials in current use. The latter materials are walled off by fibrouscapsule and are therefore less desirable for soft tissue reconstruction.At later time points (4 weeks), cellular ingrowth is even more apparentwith the appearance of vacuolar areas that may represent nascentadipocyte differentiation.

Example 10: Heparanized Formulation

A composite formulation has also been prepared with heparin conjugatedto the hyaluronic acid. This formulation was tested in vivo identicallyto the preformed scaffold above. The tissue was harvested (n=3) at 7days, 14 days, 30 days, and 90 days. Many relevant growth factors haveheparin-binding domains, such as bFGF, PDGF, and VEGF. The conjugatedheparin can serve two purposes; firstly, it can bind many of theendogenous growth factors that will be present at the injection site andserve as a local reservoir and attractive cue to the regeneratingtissue. Secondly, the heparinized composite can be used to pre-load thescaffold with growth factors to better potentiate regeneration. Theheparinized scaffolds showed enhanced angiogenesis at 7 and 14 days ascompared to the unheparinized composite scaffolds, but similar resultsat 30 and 90 days.

Example 11: Injectable Formulation

The hydrogel-nanofiber composite has also been formulated into aninjectable variant. 200 μL of the same composition as used for thepre-formed composite used in vivo (5 mg/mL thiolated HA, 5 mg/mL PEG-DA,12.5 mg/mL fibers) was mixed and allowed to partially set in the syringefor 8-10 min. At this time, the composite is a viscous, flowable liquidthat can be injected through a surgical needle (FIG. 20). Once injected,the composite maintains its shape when inverted and is non-dispersive,shape-maintaining and non/low-swelling when submersed in water. To testfor biocompatibility of the injectable composite, the suspension is theninjected into the inguinal fat pad of the rat through a 21-gauge needle.The tissue was then harvested (n=3) at 7 days, 14 days, 30 days, and 90days and analyzed identically to the previous examples. The compositedemonstrated extensive cellular remodeling at 30 days while maintainingvolume and without causing fibrotic encapsulation. Early-stageadipocytes can clearly be seen developing within the composite material.

Example 12: Use of Nanofiber Sheets to Remove the Existing WovenPolypropylene Component of Existing Surgical Meshes Altogether

In certain of the above examples, nanofiber composite technology wasused to augment and improve existing surgical mesh materials.Alternatively, the nanofiber sheets prepared during the describedprocesses could also be used to remove the existing woven polypropylenecomponent of existing surgical meshes altogether. Instead of wovenpolypropylene filaments, it was reasoned that non-woven nanofiber meshescould provide sheet structure and integrity for the surgical mesh insome applications. These could be comprised of aligned nanofiber sheetor random fiber sheet.

As an example, a composite nanofiber surgical mesh was prepared using acomposite gel formulation with 5.4 mg/mL of thiolated hyaluronic acid(220 KDa, thiolation degree of 25%) and 5.4 mg/mL of PEG-diacrylate(PEG-DA) with 10 mg/mL of functionalized dispersed nanofibers. Thecomposite gel was formed as an interpenetrating hydrogel networkimpregnating a nonwoven mesh of electrospun PCL nanofibers (taking theplace of the polypropylene filaments). The nanofiber mesh had previouslybeen functionalized to include maleimide groups on the nanofibersurface, enabling the fibers to bond directly with the surroundinghydrogel, enabling a strong formulation that is resistant todelamination at the fiber-hydrogel interface. The 1×2 cm rectangles ofnanofiber mesh were placed into the bottom of 2.5×4.5 cm Teflon molds,two per mold. An aliquot of 500 μL of composite was pipetted into eachmold (for both meshes), then a piece of plastic was placed over themeshes, and pressed down to spread out the composite. The meshes wereallowed to gel overnight in 37° C. incubator. The gelled meshes werethen lyophilized as a final product. The mesh could then be rehydratedprior to use.

This composite formulation could be used in combination with thepolypropylene mesh or nanofiber mesh could alternatively be usedindividually. The sheet was less rigid than those of the above examples,but possessed the handleability and strength characteristics appropriatefor other applications, such as wound dressings or dura repair (see FIG.22).

Example 13: Direct Surface Plasma Treatment of Woven Polypropylene orNon-Woven Microfiber Mesh Grafting with Poly(Acrylic Acid) Chains

In an alternative configuration, the woven polypropylene or non-wovenmicrofiber mesh can be directly surface plasma treated and grafted withpoly(acrylic acid) chains using the method described in PCT/US15/45494.The microfiber sheet can be used to replace the nanofiber mesh used inabove Example 12. Such a microfiber sheet can be used to form anintegrated, crosslinked network structure, as shown in FIG. 24.

As an example, a composite surgical microfiber mesh was prepared using acomposite gel formulation with 5.4 mg/mL of thiolated hyaluronic acid(220 KDa, thiolation degree of 25%) and 5.4 mg/mL of PEG-diacrylate(PEG-DA) with the appropriate amount of functionalized microfiber sheet(Ethicon Prolene Soft, Product code SPMH). The microfiber mesh hadpreviously been functionalized to include maleimide groups on the fibersurface, enabling the fibers to bond directly with the surroundinghydrogel. The composite gel was formed and crosslinked with thefunctional groups on microfibers. The 1×4 cm rectangles of Prolene Softfiber mesh were placed into the bottom of the 2.5×4.5 cm Teflon molds,two per mold. An aliquot of 500 μL of thiolated HA and PEG-DA mixture(described above) was pipetted into each mold immediately after theywere mixed. The meshes were allowed to gel overnight in 37° C.incubator. The gelled meshes were then lyophilized as a final product.The mesh could then be rehydrated prior to use (see FIG. 25).

Example 14: In Vivo Biocompatiability and Tissue Integration of theComposite Surgical Meshes

Six to eight week old Sprague-Dawley rats were randomly grouped. A 3-cmlongitudinal abdominal midline incision was marked and an incision wasmade through skin, down to the level of the abdominal wall musculature.Suprafacial dissection was continued ˜2 cm bilaterally. The 1×2 cmsurgical mesh (polypropylene, polypropylene and composite gel, ornanofiber and composite gel) was implanted in the subcutaneous planeover the right abdomen of the rat. The mesh was sutured in place, in anon-lay fashion using a 4-0 Vicryl suture. The skin was closed withinterrupted 4-0 Vicryl sutures. The left abdomen of the same ratunderwent a sham surgery. Three rats each were implanted with theunmodified Prolene surgical mesh and the composite surgical mesh withnanofiber-hydrogel network. The rats were euthanized on Days 3 and 14;and the tissue samples were explanted via en bloc resection of theabdominal wall and implanted mesh. FIGS. 26-28 show the design andoutcomes of such surgical placements.

Example 15: Discussion of Examples 5-14

Hydrogels have been widely studied as a filler material for regenerationof tissue defects due to its 3D hydrated environment and high porosity,which facilitate cell migration. However, hydrogels have proven to bepoor substitutes for volumetric defects, because the relatively weakmechanical properties of the hydrogel are insufficient to maintain itsvolume for the entire period of tissue regeneration, as the hydrogel canbe easily degraded and collapsed by body fluids and internal andexternal stresses. To improve the mechanical properties of the hydrogel,the main strategies in the field have been to (i) increase theconcentration of hydrogel precursors, (ii) increase the density of thecrosslinking network inside the hydrogel, and to (iii) introducereinforcing materials such as by embedding hydroxyapatite particles orlaminating with fiber sheets. [Mater Chem Physics, 2008, 107, 364-369,Biomaterials 2006, 27, 505-518, Acta Biomaterialia 2010, 6, 1992-2002].Unfortunately, these very strengthening strategies inherently reducedthe average pore size and porosity of the resulting hydrogel, preventingcells from being able to migrate into these hydrogels. Therefore, it wassought to strengthen hydrogels by a new mechanism that would stillretain the high porosity than allows for rapid cellular infiltration. Acomposite material was designed by introducing functionalized nanofibersthat could strengthen the overall hydrogel composite while leaving thehydrogel phase largely intact, including porosity. The resultingfiber-hydrogel composite improves upon previous soft tissue compositesbecause of two key components. Firstly, the nanofibers needed to beuniformly dispersed at a high loading level within the hydrogel toachieve isotropic strengthening. The tissue-engineering field hasgenerally utilized electrospun nanofibers as flat sheets or mats offibers. These are then typically made into composites by impregnatingthe mats with a hydrogel precursor solution.

This greatly constrains the dispersion of the nanofibers throughout thehydrogel and limits the geometry of the composites to 2D sheets ortubes. While these geometries are useful for certain applications suchas nerve repair or wound dressings, they are poor choices for repairingvolumetric defects. By cryomilling the fiber sheets, it was possible toreduce the average fiber length to the sufficiently short length thatallow them to remain in suspension in aqueous solutions. Thus, thesamples were then easily pipetted into hydrogel precursor solutions,creating a uniform dispersion of nanofiber fragments throughout thehydrogel volume before gelation. The solution can then be directly usedas an injectable formulation, or added to molds to form scaffold gels ofany arbitrary geometry, unlike the limited planar geometry of mostelectrospun nanofiber meshes. The composite structure of dispersedfibers within a hydrogel also recapitulates the fibrous architecture ofthe extracellular matrix (FIG. 6G), providing adhesion sites that mayaide cell migration within the composite.

Secondly, simply dispersing the nanofibers within the hydrogel isinsufficient to form a strong composite. These data indicated thatmerely including the nanofibers themselves provided very littleimprovement in the elastic modulus of the composite, with improvementsoccurring only when interfacial bonding was introduced. The interfacialbonding is necessary because without forming a strong linkage betweenthe hydrogel and fiber components, the water and hydrogel componentscould slide past the fiber components without transferring the loads tothe stiffer material. Furthermore, the interface between such disparatematerials could lead to delamination and failure in the composite.Further, PCL's initial hydrophobicity makes it difficult to disperse inaqueous solutions, as the fibers preferentially clump together and formclots that fall out of suspension. Plasma treatment and subsequentfunctionalization with carboxylic acid groups and amine groups greatlyincreases the hydrophilicity of the fibers and allows dispersion. Thedramatic increase in mechanical properties only occurred wheninterfacial bonding occurred between the maleimide groups on the fibersurfaces and the thiol groups on the hyaluronic acid molecules. Thiscovalent-strength bonding transfers loads more efficiently to the fibersduring compression or tension, leading to a stiffer, stronger material.Moreover, the composites show a strong trend of increasing elasticmoduli with increasing maleimide density, emphasizing its primacy in thestrengthening mechanism, as well as the tunable nature of thereinforcement.

In this study, it was identified that it was possible to tune themechanical properties of the fiber-hydrogel composite by variousfactors, including the total surface area of fibers, the density of thefunctional maleimide groups on the fiber surface, and the amount offibers loaded into the hydrogel. Firstly, composites with smallerdiameter of fibers showed a higher compressive and shear storage modulusthan those of composites with bigger diameter of fibers (FIG. 7A andFIG. 8C). Similarly, in the literature, a single ultra-highmolecular-weight polyethylene (UHMWPE) fiber (25 μm), which was plasmaactivated using glutaraldehyde, showed approximately 2.36-fold increaseof interfacial shear strength in a poly(vinyl alcohol) hydrogel comparedto it of a UHMWPE fiber bundle of 60 [Acta Biomaterialia 2014, 10,3581-3589]. Therefore, it is possible that decreasing the fiber diameterand thus increasing the fiber specific surface area may be an effectivein improving the mechanical properties of the composite. However, eachfiber group had a slightly different MAL surface density on the fibers(approximately 10-15 nmol/mg), so the effect of surface area of fibersalone cannot definitively be determined. Hence, secondly, compositeswere fabricated with the same diameter fibers, but with various MALsurface densities on the fibers (FIG. 8). The compression and shearstorage moduli of composites were increased with increasing MAL surfacedensity on the fibers. It was confirmed that a composite without theinterfacial bonding showed only a slight enhancement of its compressivemodulus (FIG. 7) by using fibers modified through the PAA step (carboxylgroups on fibers), but not the further MAL conjugation steps. Theimportance of the interfacial bonding was additionally confirmed byquenching the MAL groups on the fibers with cysteine prior to gelation.The cysteine conjugates to the maleimide group and prevents interfacialbonding between the fibers and hydrogel, which allows us to isolate justthe effect of interfacial bonding, since the fibers were otherwiseprocessed identically to the interfacial-bonding groups. Interestingly,the mechanical properties of the composites with the MAL-quenched fiberswere dramatically diminished (FIG. 7A and FIG. 8B), with theMAL-quenched fiber group showing a lower compressive modulus than thatof HA hydrogel-alone when the concentration of HA was 10 mg/ml (FIG. 7).It is possible that the MAL-quenched fibers weakened the overallcomposite by delaminating easily at the interface of the fibers andhydrogel, as is seen in previous studies [Acta Biomaterialia 2014, 10,3581-3589]. Also, the fibers without functional groups may be acting asan alien substance that inhibits gelation compared to a pure hydrogelcomposed of one component or without any alien substance during gelation[JMC B 2015, DOI: 10.1039/C3TB21830A, Journal of Biomedical MaterialsResearch Part A 2010, 95 (2), 564-573]. Furthermore, a significantcorrelation between shear storage modulus and the density of theinterfacial bonding by composites with various MAL surface densities wasverified (FIG. 8C). These studies provide strong evidence that themechanical properties of hydrogel could be reinforced and tuned by theinterfacial bonding. Thirdly, the shear storage modulus of thecomposites was enhanced with an increasing weight ratio of fibers tohydrogel (FIG. 9). Thereby, it was confirmed that the weight ratio wasanother variable that can be used to tune the mechanical properties of afiber-hydrogel composite. However, here, it was confirmed that withincreasing fiber loading, the shear storage modulus increases began tolevel off and even slightly decreased above 0.6 of the weight ratio. Onepossibility for this saturation effect may be that the density ofinterfacial bonding of a composite was diminished by how the excessfibers with MAL reacted with a large fraction of the thiol groups of HA,preventing them from reacting with the PEGDA for gelation. Consideringthat the highest shear storage modulus of the HA hydrogel was obtainedwith an equimolar amount of each functional group of HA-SH and PEGDA aswell as the decreasing shear storage modulus with excess amounts ofeither HA-SH or DA (FIG. 14A), the excess MAL on the fibers with theincreasing amount of fibers could disrupt the SH-to-DA bonding inside acomposite.

Generally, implanted biomaterials have to withstand numerous internaland external stresses during regeneration of the tissue defect. Althoughthe stress is not severe and continuous, stress resistance tests wereperformed under a repeating condition and a high frequency (10 Hz) tomimic such stresses (FIG. 10 and FIG. 8). Both the HA hydrogel andfiber-HA hydrogel composite withstood without any damage or reduction oftheir mechanical strength during repeating compressive strain.Noticeably, composites with the interfacial bonding retained their shearstorage modulus at 10 Hz of frequency, whereas the shear storage modulusof the HA hydrogel and the composite without the interfacial bondingwere diminished at 10 Hz. This trend indicates that the interfacialbonding with the dispersed fibers is crucial to the reinforcement of thecomposite's mechanical properties. In addition, the fiber-HA compositesmaintained their dimensions and their Young's moduli after beingsubjected to lyophilization and subsequent rehydration, while theHA-alone gel shrank substantially under the same process (FIG. 6C andFIG. 10). This shape, volume, and stiffness maintenance afterdehydration and rehydration is an important feature for clinicaltranslation of this technology, as having a lyophilized form of thecomposite would make it easier to sterilize and store the commercialproduct.

For soft tissue reconstruction, the ideal implanted scaffold wouldimmediately fill the defect void, but would also serve as a substratefor the body's own cells to grow into the scaffold, proliferate anddifferentiate into the proper tissue phenotype, eventually replacing theartificial scaffold with normal, healthy tissue. Thus, it is criticallyimportant that relevant cells would be able to migrate within thehydrogel or composite scaffold To determine the potential for relevantcell types to migrate within the scaffolds, hASC spheroids were seededinside HA hydrogels and fiber-HA hydrogel composites and evaluated theircell migration. The hASCs could not migrate inside the HA hydrogel-alonebecause the HA hydrogel was too soft to serve the traction forces forcell migration (FIG. 11A) [Biomaterials 2015, 42, 134-143].Interestingly, although shear storage modulus of the composites wassimilar to that of the HA hydrogel, the hASCs were able to significantlymigrate away from a spheroid inside the composites (FIG. 11). Onehypothesis is that the fibers inside a composite may be providingadhesion sites to guide cell migration similarly to the fibrilcomponents of the native ECM of adipose tissue. It was previouslydemonstrated that aligned and random fibers could be a critical factorfor cell adhesion, proliferation, differentiation, and migration invarious cell types [Biomaterials 2005, 26, 2537-2547/2006, 27,6043-6051/2009, 30, 556-564/2010, 31, 9031-9039, Acta Biomaterialia2013, 9, 7727-7736]. Especially, it was observed that cells recognizedfibers as a guide matrix, as their cytoskeletons aligned with andfollowed along the underlying fibers [Biomaterials 2006, 30,6043-6051/2009, 30, 556-564]. However, the diameter of the fibers insidethe composites did not affect the migrating cells, as they migratedrobustly in composites with either 1000-nm or 286-nm nanofibers (FIG.19).

The porosity and cell migration effects seen in benchtop testing and invitro cell culture translated into profound differences during in vivotesting of the composites. The hydrogels formulated to fat-mimicking 2kPa stiffness without fibers had a porosity too low for cellularinfiltration. The cellular response was to wall off the hydrogel with athick layer of collagen, with the lack of infiltration or remodelingtypical of a foreign body response. The nanofiber-hydrogel composite,however, had sufficient porosity to facilitate cellular ingrowth,vascularization, and cellular remodeling without the foreign bodyresponse. This offers the prospective of permanently filling thevolumetric defect in the body with what will ultimately be the body'sown tissue. The results were even more pronounced in the injectableformulation, which can form a tighter interface with the host tissue andshowed signs of robust adipogenesis.

Conclusion:

The dispersion of functionalized nanofibers within a hydrogel forms acomposite structure with the combined strengths of the two components.The interfacial bonding between the nanofibers and the hydrogelcomponents is critical to making a strong composite, while maintaininghigh porosity and pore size to facilitate tissue and cell ingrowth. Theresulting composite properties can be easily tuned by varying the fiberdiameter, fiber loading level, maleimide density level, and the loadinglevels of the hydrogel components. This allows for lower crosslinkingand higher porosity at a targeted overall stiffness, increasing cellularinfiltration and subsequent tissue remodeling. The fibers themselves mayalso directly improve cellular migration by providing adhesion sitessimilar to that seen in the native ECM. The resulting composite implantcan be tuned to match the stiffness of native fat tissue, yet remainpermeability for cellular infiltration and remodeling. This novelcomposite is strong enough to immediately fill a volumetric defect ofany arbitrary shape. The composite implant then serves as a permissivescaffold for the body's own cells to infiltrate into the composite, formblood vessels, and differentiate into cells like adipocytes. Thescaffold will be slowly degraded away during tissue remodeling, untilthe initial defect void has been replaced fully by normal, healthytissue. The composite structure has great potential for reconstructiveand aesthetic surgery potential.

Example 16. Production and Use of Medical Devices

Synthetic and biologic meshes have wide applicability in generalsurgery, reconstructive surgery, neurosurgery, urology, gynecologicsurgery, orthopedic surgery, and aesthetic surgery. These meshes areused to reinforce or replace tissues throughout the body. A majorlimitation of these meshes is their poor integration with the bodyleading to foreign body reactions, seroma formation, and infection amongother complications. Provided are a nanofiber-hydrogel compositematerial that has greatly improved integration with the body's tissuecompared to existing synthetic meshes. This material is produced andutilized as flat sheets to serve as a mesh for tissue reinforcement andreplacement. It is also useful as an adjunct to conventional meshes toimprove their biocompatibility and integration potential.

The scaffold complexes of the invention are incorporated into surgicaldevices by association with a surgical mesh material. For example, asurgical device contains i) a “laminar” scaffold complex comprising apolymeric fiber having a mean diameter of from about 100 nm to about8000 nm operably linked to a hydrogel material; and ii) a surgical meshmaterial. As used herein, a “laminar” scaffold complex is generallyplanar and flexible, and having a sufficient surface are with which itcan associate with the surgical mesh. Exemplary surgical meshes includepolyglactin 910, polypropylene, polyglycolic acid,polytetrafluoroethylene (ePTFE), polypropylene withpolyglactin-absorbable, polypropylene with poliglecaprone 25,polypropylene with cellulose, macroporous polypropylene with ePTFE. Suchmedical devices are of a size and shape such that they can be readilyimplanted (i.e., inserted) into a human subject at or proximal to thesite of a tissue defect (e.g., at the site of surgical treatment). Thescaffold complex is covalently or non-covalently associated with thesurgical mesh.

EQUIVALENTS

It is understood that the detailed examples and embodiments describedherein are given by way of example for illustrative purposes only, andare in no way considered to be limiting to the invention. Variousmodifications or changes in light thereof will be suggested to personsskilled in the art and are included within the spirit and purview ofthis application and are considered within the scope of the appendedclaims. For example, the relative quantities of the ingredients may bevaried to optimize the desired effects, additional ingredients may beadded, and/or similar ingredients may be substituted for one or more ofthe ingredients described. Additional advantageous features andfunctionalities associated with the systems, methods, and processes ofthe present invention will be apparent from the appended claims.Moreover, those skilled in the art will recognize, or be able toascertain using no more than routine experimentation, many equivalentsto the specific embodiments of the invention described herein. Suchequivalents are intended to be encompassed by the following claims.

What is claimed is:
 1. A surgical scaffold device comprising a laminarscaffold complex, comprising: polymeric fibers having a mean diameter offrom 100 nm to 8000 nm covalently linked to a hydrogel material,wherein: a) the polymeric fibers are dispersed throughout the hydrogelmaterial; b) the scaffold complex comprises a reacted crosslinkingmoiety; c) the scaffold complex comprises a plurality of pores presenton or within a surface of the scaffold complex, wherein the pores arepresent at a concentration of at least 50 pores per cm², and wherein atleast 80% of the pores have an average pore diameter of at least 5microns; d) the device is configured in a geometry suitable forimplantation in an organ or tissue selected from skin, fascia, pleura,dura, pericardium, paratenon, periosteum, perineurium, blood vesselwall, and lymphatic wall; and e) the device is configured as a sheethaving a first dimension and a second dimension independently at leastfive times as great in length as a third dimension.
 2. The surgicaldevice of claim 1, wherein the hydrogel material comprises afunctionalized hyaluronic acid and the polymeric fibers comprises afunctionalized polycaprolactone.
 3. The surgical device of claim 1,wherein the laminar scaffold complex is isotropically reinforced.
 4. Thesurgical device of claim 1, wherein the polymeric fibers comprisefunctional groups and the surface density of the functional groups onthe polymeric fibers is from 10 nmole/mg of fibers to 160 nmole/mg ofthe fibers.
 5. The surgical device of claim 1, wherein the reactedcrosslinking moiety comprises a reacted maleimide group.
 6. The deviceof claim 1, configured as a sheet having a first dimension and a seconddimension independently at least five times as great in length as athird dimension.
 7. The device of claim 6, wherein the third dimensionis less than 1 cm.
 8. The device of claim 6, wherein the third dimensionis less than 5 mm.
 9. The device of claim 6, wherein the third dimensionis less than 1 mm.
 10. The device of claim 6, wherein the firstdimension and second dimension are independently greater than 1 cm. 11.The device of claim 7, wherein the first dimension and second dimensionare independently greater than 1 cm.
 12. The device of claim 1,configured to be subdermally implantable in a human subject in needthereof.
 13. A medical device for use in a subject undergoing a surgicalprocedure, comprising the device of claim 1 in an amount effective toprovide for the reinforcement and/or regeneration of one or more tissuesincluding skin, fascia, pleura, dura, pericardium, paratenon,periosteum, perineurium, blood vessel wall, lymphatic wall whenadministered to the subject, or the reinforcement or coverage of amedical implant when administered to the subject.
 14. A surgical devicecomprising: i) a laminar scaffold complex comprising polymeric fibershaving a mean diameter of from 100 nm to 8000 nm covalently linked to ahydrogel material; wherein the polymeric fibers are dispersed throughoutthe hydrogel material, the scaffold complex comprises a reactedcrosslinking moiety, the scaffold complex comprises a plurality of porespresent on or within a surface of the scaffold complex, wherein thepores are present at a concentration of at least 50 pores per cm², andwherein at least 80% of the pores have an average pore diameter of atleast 5 microns; ii) a surgical mesh material; wherein the surgical meshmaterial is covalently linked to the laminar scaffold complex, andwherein the surgical mesh material is disposed over all or a portion ofthe laminar scaffold complex.
 15. The surgical device of claim 14,wherein the polymeric fibers comprise a synthetic polymeric materialcomprising a poly(lactic-co-glycolic acid), poly(lactic acid), and/or apolycaprolactone, or a combination thereof; or a biological polymericmaterial selected from the group consisting of a silk, collagen,elastin, hyaluronic acid, chitosan, or a combination thereof.
 16. Thesurgical device of claim 14, wherein the hydrogel material comprises apoly(ethylene glycol), a collagen, a dextran, an elastin, an alginate, ahyaluronic acid, a poly(vinyl alcohol), or a combination thereof. 17.The surgical device of claim 14, wherein the surgical mesh materialcomprises fibers selected from the group of polycaprolactone,polypropylene, polyglycolic acid, polytetrafluoroethylene (ePTFE),polyglactin, poliglecaprone, cellulose, or a combination thereof. 18.The surgical device of claim 14, wherein the reacted crosslinking moietycomprises a reacted maleimide group.
 19. The surgical device of claim14, configured to be subdermally implantable in a human subject in needof.
 20. The surgical device of claim 14, comprising a first laminarscaffold complex and a second laminar scaffold complex, wherein thesurgical mesh material is disposed between the first laminar scaffoldcomplex and the second laminar scaffold complex.
 21. The surgical deviceof claim 14, comprising a first laminar scaffold complex and a secondlaminar scaffold complex, wherein the surgical mesh material isinterposed between the first laminar scaffold complex and the secondlaminar scaffold complex.
 22. The surgical device of claim 14, furthercomprising a human cell or cell extract.
 23. The surgical device ofclaim 14, wherein the hydrogel material is present in the complex in afunctional network.
 24. The surgical device of claim 14, furthercomprising a therapeutic agent selected from a cell, a small molecule, anucleic acid, and a polypeptide.
 25. The surgical device of claim 14,wherein the polymeric fiber is arranged in a sheet.